DEVELOPING AN ANATOMICAL SIMULATOR FOR EARLY STAGE USABILITY TESTING OF ENDOVASCULAR DEVICES KEVIN JEREMIAH O SULLIVAN, B.SC., AIES, CSWP.

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1 DEVELOPING AN ANATOMICAL SIMULATOR FOR EARLY STAGE USABILITY TESTING OF ENDOVASCULAR DEVICES KEVIN JEREMIAH O SULLIVAN, B.SC., AIES, CSWP. Submitted for the degree of Master of Science (M.Sc.) Design Factors Research Group University of Limerick Ireland. Research Supervisors: Dr. Leonard O Sullivan, Senior Lecturer in Ergonomics Louise Kiernan, Lecturer in Product Design & Technology Date Submitted: May 2016

2 Declaration I hereby declare that my submission is the result of my own work, and as a whole is not substantially the same as any that I have previously made or am currently making, whether in published or unpublished form, for a degree, diploma, or similar qualification at any university or similar institution. Kevin J. O Sullivan B.Sc., CSWP, AIES. i

3 Acknowledgment I would like to thank my academic supervisors Dr. Leonard O Sullivan and Louise Kiernan as well as Design Partners and the Irish Research Council for Science, Engineering and Technology (IRCSET) for funding this research. ii

4 Abstract Adverse medical events resulting in serious harm occur in approx. 1-4% of all hospital admissions (Sarker and Vincent 2005). Endovascular devices encompasses a wide range of medical devices designed to treat diseases using a minimally invasive approach. These devices range from coronary stents to full aortic valve replacement and are delivered through the vascular system. Usability engineering is a sub specialty of ergonomics concerned with the development of devices that are fit for use (Kramme et al. 2011). Usability Testing (UT) of medical devices is an integral part of the regulatory requirements for device approval in both the United States and the European Union. UT champions the use of simulated use testing, particularly at the early stage and validation stage of the design process (FDA 2011). Simulated UT is a safe and ethical way of eliciting user needs, testing device prototypes, and validating design solutions in order to reduce the likelihood of an adverse event occurring. During UT, there is an onus on the tester to provide operators with a sufficient level of fidelity so as to cause a suspension of disbelief during the task (Halamek et al. 2000); that is, that the operator thinks and feels like they are performing in the real environment. Currently, UT of endovascular medical devices is undertaken in pulsatile flow rigs with hollow silicone anatomical models and direct visualisation of devices in the model. There are several limitations with the use of these models such as the lack of anatomical realism, manufacturing methods and portability. These omissions can negatively impact the reaction of endovascular devices in the models, resulting in flawed test results. The aim of the current research was to develop an anatomical simulator for early stage usability testing of endovascular devices. A protocol was developed to digitally segment compound anatomical models that include all structures encountered in clinical practice such as thrombus and calcifications. These digital models were modified to include standardised mounts for the usability test bed. New methods of physically reproducing compound models using multi material 3D printing were explored, as the use of the lost wax casting method is inappropriate for these complex models. The 3D printed compound anatomical models were integrated into a newly developed portable pulsatile flow simulator. The simulator includes real time haemodynamic monitoring and a simulated fluoroscopy imaging system. The result of the current work was the successful segmentation of compound anatomical models and the subsequent manufacture using multi material 3D printing. These 3D printed compound anatomical models were then tested in a newly developed portable pulsatile flow simulator that includes an on board simulated imaging system and real time haemodynamic monitoring of the model. iii

5 List of abbreviations AAA AS BA BPM CAD CCD CE CNC CO COPD CPB CTA DAQ DICOM DPI EC EMG EVAR FDA HDPE HF HFE HU IFU ILT IOM ITU LBBB LVOT MVE OR PAD POC PPI R&D ROI SAVR SEIPS SS TAVI Tg UCD UE US UT UV VGA VI Abdominal Aortic Aneurysm Aortic Stenosis Balloon Angioplasty Beats per Minute Computer Aided Design Charge Coupled Device Conformity European Computer Numerical Controlled Cardiac Output Chronic Obstructive Pulmonary Disease Cardio Pulmonary Bypass Machine Computed Tomography Angiography Data Acquisition Digital Imaging and Communications in Medicine Dots Per Inch European Commission Electromyography Endovascular Aneurysm Repair Food and Drug Administration High Density Polyethylene Human Factors Human Factors Engineering Hounsfield Units Instructions For Use Intra Luminal Thrombus Institute Of Medicine Intensive Care Unite Left Branch Bundle Block Left Ventricular Outflow Track Maximum Voluntary Exertion Operating Room Peripheral Artery Disease Proof Of Concept Permanent Pacemaker Implantation Research & Development Region Of Interest Surgical Aortic Valve Replacement System Engineering Initiative for Patient Safety Stainless Steel Transcatheter Aortic Valve Implantation Glass Transition Temperature User Centred Design Usability Engineering United States Usability Testing Ultra Violet Video Graphic Array Virtual Instrument iv

6 Table of contents DECLARATION... I ACKNOWLEDGMENT... II ABSTRACT... III LIST OF ABBREVIATIONS... IV TABLE OF CONTENTS... V LIST OF FIGURES... IX LIST OF TABLES... XII PUBLISHED WORK... XIII INTRODUCTION BACKGROUND THESIS STRUCTURE...3 CHAPTER 2: LITERATURE REVIEW ADVERSE EVENTS IN MEDICINE HUMAN FACTORS AND ADVERSE EVENTS IN MEDICINE ADVERSE EVENTS IN EVAR ADVERSE EVENTS IN TAVI HUMAN FACTORS THEORETICAL MODELS SYSTEMS ENGINEERING INITIATIVE FOR PATIENT SAFETY (SEIPS) SEIPS CONFIGURATION ENGAGEMENT PROFESSIONAL WORK USE OF THE SEIPS MODEL IN THIS RESEARCH v

7 2.2. ROLE OF MEDICAL DEVICE DESIGN IN PREVENTING ERRORS MEDICAL DEVICE DESIGN CONTROLS APPLICATION OF DESIGN CONTROLS USER CENTRED DESIGN OF MEDICAL DEVICES USABILITY USABILITY TESTING OF MEDICAL DEVICES DURING DESIGN FORMATIVE TESTING SUMMATIVE (VALIDATION) TESTING SIMULATION IN USABILITY TESTING OF MEDICAL DEVICES SUSPENSION OF DISBELIEF FIDELITY IN SIMULATION SIMULATORS USED IN MEDICAL CONTEXTS SILICONE ANATOMICAL MODELS PROCESS OF FORMING SILICONE ANATOMICAL MODELS COMMERCIAL SILICONE MODELS LIMITATIONS OF HOLLOW SILICONE ANATOMICAL MODELS IMAGING IN ENDOVASCULAR SIMULATION DEVELOPMENT OF USABILITY SIMULATORS FOR AAA AND TAVI ABDOMINAL AORTIC ANEURYSM (AAA) AORTIC STENOSIS DIFFICULTIES WITH ENDOVASCULAR DELIVERY SYSTEM USE SUMMARY OF LITERATURE REVIEW RESEARCH OBJECTIVES CHAPTER 3: SEGMENTATION OF COMPOUND ANATOMICAL MODELS INTRODUCTION AND RESEARCH OBJECTIVE PROCESS FOR SEGMENTING ANATOMICAL MODELS vi

8 3.3. SEGMENTATION OF ANATOMICAL STRUCTURE TO CREATE COMPOUND MODELS SEGMENTATION APPROACH FOR AAA AND TAVI SEGMENTED MODEL MODIFICATION FOR USE IN PULSATILE SIMULATORS SUMMARY OF OUTCOMES FROM THIS CHAPTER CHAPTER 4: 3D PRINTING OF ANATOMICAL MODELS FOR USABILITY TESTING INTRODUCTION AND RESEARCH OBJECTIVE D PRINTING TECHNOLOGIES FOR ANATOMICAL MODELS STEREOLITHOGRAPHY (SLA) SELECTED LASER SINTERING (SLS) FUSED DEPOSITION MODELLING (FDM) INKJET PRINTING LAMINATED OBJECT MANUFACTURING (LOM) POLYJET PRINTING TECHNOLOGY POLYJET PRINTING USING OBJET CONNEX MATERIALS PROCESS OF 3D PRINTING COMPOUND ANATOMICAL MODELS INITIAL EXPERIENCE WITH COMPOUND ANATOMICAL MODELS RESULTS OF 3D PRINTED COMPOUND MODELS SUMMARY OF OUTCOMES FROM THIS CHAPTER CHAPTER 5: DEVELOPMENT OF A USABILITY TEST BED FOR 3D PRINTED COMPOUND ANATOMICAL MODELS INTRODUCTION AND RESEARCH OBJECTIVE MECHANICAL DEVELOPMENT PULSATILE FLOW PRESSURE vii

9 VISCOSITY TEMPERATURE SIMULATED IMAGING LABVIEW INTERFACE ELECTRONIC CONTROLS HOUSING AND PORTABILITY RESULTS OF PORTABLE USABILITY TEST BED PROOF OF CONCEPT USABILITY TEST USING THE COMPLETED SIMULATOR SUMMARY OF OUTCOMES FROM THIS CHAPTER CHAPTER 6: DISCUSSION OBJECTIVE ONE: SEGMENTING COMPOUND MODELS OBJECTIVE TWO: MODIFYING ANATOMICAL MODELS OBJECTIVE THREE: 3D PRINTING COMPOUND ANATOMICAL MODELS OBJECTIVE FOUR: DEVELOPING A PORTABLE USABILITY TEST BED OBJECTIVE FIVE: SIMULATED IMAGING AND MONITORING CONTRIBUTION OF RESEARCH LIMITATIONS FUTURE WORK CHAPTER 7: CONCLUSIONS REFERENCES... I APPENDIX I: ARDUINO CODE... XVI APPENDIX II: LABVIEW CODE... XVII APPENDIX III: TECHNICAL DRAWINGS OF SIMULATOR... XVIII APPENDIX IV: FDA 21 CFR DESIGN CONTROLS... XIX viii

10 List of figures Figure 1: Swiss Cheese Model of Accident [adverse event] Causation (Reason 2000) Figure 2: Human Information Processing Model (Hollands and Wickens 1999) Figure 3: Model of Factors Influencing Surgical Errors (Sarker and Vincent 2005) Figure 4: Application of design controls to Waterfall Design Process (Medical Devices Bureau, Health Canada) Figure 5: Relationship between Risk Management and Usability Engineering Standards Figure 6: Integrating Human Factors into the Design Process Figure 7: Dimensions of Fidelity and cost in simulation Figure 8: Top, AAA simulator (Chong et al. 1998). Bottom, Pulsatile Flow Loop by CABER, University of Limerick Figure 9: Addition of ILT to AAA models (Ene et al. 2011) Figure 10: Silicone AAA model (A-S-A005+) by Elastrat, Geneva, Switzerland Figure 11: Fluoroscopy image showing a Sapien aortic valve before and during balloon inflation (Cheung and Lichtenstein 2012) Figure 12: Illustration of infrarenal AAA in the body (Society of Interventional Radiology, 2010) Figure 13: Cross section of the aorta, from Holzapfel et al. (2000) Figure 14: Management plan for treatment of AAA, from Sakalihasan et al. (2005) Figure 15: schematic view of aorta and aortic valve Figure 16: Heavily Calcified aortic valve ix

11 Figure 17: 30 day mortality rates in TAVI. First 135 patients and second 135 (Gurvitch et al. 2011) Figure 18: Medtronic CoreValve Evolute Delivery System Figure 19: The process chain of Medical Image to 3D Printed Model (Rengier et al. 2010) Figure 20: User interface of the Mimics Program Figure 21: Sagittal View of AAA Patient Figure 22: Sagittal view with True Lumen Mask Figure 23: Sagittal view with Aneurysm Sac masked Figure 24: Sagittal View with Calcifications masked Figure 25: Sagittal View with all parts combined Figure 26: 3D Reconstruction of an AAA shown in the context of the 2D CT scan Figure 27: Completed Segmentation of AAA, Lumen, Calcifications and ILT with cross Sectional view and Location indicator Figure 30: Sagittal view with True Lumen Masked Figure 30: Sagittal View with 3D reconstruction of True Lumen Figure 30: Sagittal View with Post processed Aortic model Figure 31: Top: Internal Lumen (Contrast Path) of the left ventricle and Aorta. Bottom: Hollow, 3D Printer ready file with Rigid Connection Ports Figure 32: Full colour helmets printed on an Objet Connex 3, Stratasys Ltd. 82 Figure 33: Complete Aorta printed on the Connex Figure 34: AAA model in the Object Studio Software Figure 35: Printed AAA with section removed showing Printing defect Figure 36: Cross section of AAA showing Aortic wall, Thrombus, Calcifications, and True Lumen Figure 37: Cross section of Iliac artery showing calcifications Figure 38: left, a digital view of the hollow anatomical model with rigid ports attached, right, the 3D printed result x

12 Figure 39: 3D Printed left ventricle, rigid port and silicone tubing Figure 40: Additional AAA models printed including Calcifications Figure 41: (A) Computer Model of AAA, (B) 3D Printed Model on the printer bed, (C) 3D Printed model after cleaning Figure 42: Schematic of pulsatile simulator Figure 43: Solidworks model of planned simulator Figure 44: Output to Pressure Differential, Freescale MPX5050GC7U Figure 45: Simulated fluoroscopic image on the simulator Figure 46: LabView interface of the Simulator Figure 47: Front view of the completed Simulator Figure 48: Left, Operator performing femoral artery angioplasty, Right, simulated fluoroscopy image on display during procedure Figure 49: 3D printed generic TAVI delivery system Figure 50: Interchangeable handle for generic delivery system Figure 51: peak forces generated by sex and grip type Figure 52: Individual maximum voluntary exertion plots for Pilot Test Figure 53: Installed aortic model with left ventricular input port Figure 54: Encapsulated materials Figure 55: Delamination along the Z plane (aorta) Figure 56: Picasso Haemodynamic monitor Figure 57: Simulated image using x-ray film as background (Chong et al. 1998) xi

13 List of tables Table 1: Summary of pertinent literature on TAVI adverse events Table 2: SEIPS applied to TAVI, Adapted from Holden et al. (2013) Table 3: HFE Goals of Formative Testing (FDA 2011) Table 4: Advantages of simulators (Maran and Glavin 2003) Table 5: Limitations of hollow silicone phantoms for device testing Table 6: TAVI delivery Systems Table 7: Retrograde TAVI Delivery System Task Analysis Table 8: Antegrade TAVI delivery System Task Analysis Table 9: Overview of established 3D printing techniques used in the medical arena (Rengier et al. 2010) xii

14 Published work The following is a list of papers presented, published or under review based on or associated with the work presented in this thesis: Moore, S. S., O Sullivan, K. J. and Verdecchia, F. (2015) 'Shrinking the Supply Chain for Implantable Coronary Stent Devices', Annals of biomedical engineering, O'Sullivan, K. J., O'Sullivan, L. W., Kiernan, L. and Canavan, E. (2015) 'Developing a Virtual Beating Arterial Model for early stage usability testing', Mimics World Conference. O'Sullivan, K. J., O'Sullivan, L. W., Kiernan, L. and Canavan, E. (2014) 'Integrating Human Factors Engineering Into the Design Process', Human Factors and Ergonomics Society. Curtis, E., O'Sullivan, K. J., White, E. and de Eyto, A. (2013) 'Rise of the machines: Has rapid prototyping evolved faster than the software used to create it? The limitations of computer aided design in rapid prototyping in rapid prototyping of organic forms', High Value Manufacturing. O'Sullivan, K. J., O'Sullivan, L. W., Kiernan, L. and Canavan, E. (2013a) 'A Human Factors Performance Model of Endovascular Surgery Applied to Transcatheter Aortic Valve Implantation', Liepzig Interventional Conference. O'Sullivan, K. J., O'Sullivan, L. W., Kiernan, L. and Canavan, E. (2013b) 'Modelling the effects of error producing conditions on adverse events during transcatheter aortic valve implantation', Irish Ergonomics Reveiw O'Sullivan, K. J., O'Sullivan, L. W., Kiernan, L. and Canavan, E. (2013c) 'Using biofidelic phantom organs for design stage usability testing of endovascular delivery systems', High Value Manufacturing. xiii

15 Introduction 1.1. Background The endovascular device market (excluding cerebral intervention and transcatheter aortic valve implantation) is expected to reach $26.72 billion by 2019 (Maketsandmarkets 2015). Over one million patients alone undergo cardiac catheterisation in the US annually (Mozaffarian et al. 2015). Endovascular interventions are used to treat a wide range of clinical areas including vascular, oncological, neurological, nephrology, and gastroenterology. Procedures from angiography and percutaneous coronary intervention, to cerebral aneurysm coiling, thoracic aneurysm repair, chemoembolization, and aortic valve replacements can all be accomplished using these minimally invasive methods. The procedures can be performed by various operators such as cardiologists, interventional radiologists, or vascular surgeons. Adverse events in medicine are any action that results in harm to a patient. These can include erroneous diagnostic tests (wrong test requested or wrong information provided), medication errors (wrong medication, dose or administration route), up to and including wrong side surgery (such as the removal of the wrong kidney during radical nephrectomy). Most adverse events do not result in serious harm to the patient, however, in approximately 1 to 4% of hospital admissions, serious injury does occur (Sarker and Vincent 2005). Of this cohort, 47.7% were associated with a surgical operation. It is not currently known how much device design contributes to this figure. The historical view of medicine is that healthcare professionals should perform without error (de Leval 1997). Since the 1980 s, there has been a move towards understanding what external factors lead to an adverse event. The contemporary model is that errors are to be expected as humans are fallible, 1

16 even in the best organisations. The majority of adverse events result not from the action of an isolated individual, but as the result of deficiencies in the system of which they are a part (Carayon et al. 2006, Carayon 2011). These deficiencies can be in the organisational structure (how many hours on shift), the technology available (is there access to the best of equipment), the macroeconomic climate (is there funding to provide the best care) and individual stresses (personal life) amongst others. Endovascular procedures are minimally invasive medical procedures delivered primarily through the vascular system. To perform these procedures there is a need for robust, well designed, safe and effective medical devices. To achieve this, usability testing is prescribed from the earliest inception of a medical device to the validation testing prior to market launch (FDA 1997, FDA 2011, Alley 2014, Home and FAQs 2010). Usability testing relies heavily on simulated use testing: Testing devices in anatomical models and under simulated use conditions (environmental). While there is a regulatory requirement to undertake usability testing, there are no standardised methods to do so. Currently, this testing is primarily carried out using hollow silicone anatomical models in laboratory based pulsatile flow simulators. There are however, several limitations to the current approach. To address these deficiencies, the current work was undertaken with the overall objective: To segment compound anatomical models for the specific purpose of physically recreating the models using multi material 3D printing, to modify the resulting models in such a way that they can be integrated into a portable usability test bed that includes real time haemodynamic monitoring and simulated fluoroscopic imaging. 2

17 1.2. Thesis structure Chapter Two provides an overview of the current literature pertaining to adverse events in medicine, specifically relating to AAA and TAVI. The role of human factors methodologies in identifying and addressing potential causes of adverse events is also described. A brief overview of the disease states abdominal aortic aneurism and aortic stenosis is presented including background, causes, and treatment options. The latter part of the chapter provides an overview of the importance of usability testing and the current methods and limitations for testing endovascular devices. Chapter three describes the method used to segment compound anatomical models using the Mimics software program. Four individual digital anatomical models (three AAA and one TAVI) were created from anonymised patient CT data with a view to recreating the models using a multi material 3D printer. This chapter also details how compound anatomical models can be modified in order to enhance functionality of the models such as the addition of standardised mounts for integrating into pulsatile flow simulators. Chapter four details the method used to physically recreate compound anatomical models using multi material 3D printing. An overview of 3D printing technology used in the medical industry is presented and the suitability of polyjet printing demonstrated. Initial experiments are reported and a revised segmentation protocol developed as a result of the initial findings. Chapter five documents the building of a portable usability test bed for use with the 3D printed compound anatomical models described previously. This chapter also presents the integration of real time haemodynamic monitoring as well as a simulated fluoroscopic imaging system into the test bed. A pilot usability test of a TAVI delivery system is documented at the end of the chapter. 3

18 The pilot test was conducted using electromyography to evaluate peak muscle use and force during deployment of a TAVI valve. Chapters six and seven include the discussion and conclusions respectively, alongside limitations and proposed future work. 4

19 Chapter 2: Literature review The advantages of endovascular procedures over open surgical procedures include reduced mortality, reduced hospital stays and improved recovery times for patients. However, as with all medical procedures they are not without risk of adverse events. The Food and Drug Administration (FDA) defines an adverse event as: Any undesirable experience associated with the use of a medical product in a patient. Much work has been undertaken to understand factors that can influence or even cause adverse events to occur in clinical settings. A wide body of literature has discussed the importance of considering human factors in medical contexts, but relatively few studies have specifically presented an analysis of the design of devices in an attempt to understand the links between design and effectiveness, efficiency or satisfaction (Sharples et al. 2012). There is, however, a growing body of literature highlighting the role of medical device design in preventing (or facilitating) adverse events. For the purposes of this research, two disease states that have readily adopted endovascular treatment options were selected: Endovascular Aneurysm Repair (EVAR) and Transcatheter Aortic Valve Implantation (TAVI). EVAR, an endovascular treatment for Abdominal Aortic Aneurysms (AAA), represents a mature technology (first trialled in the late 1980 s) while TAVI, a treatment for Aortic Stenosis (AS), is a rapidly advancing field still in its genesis (first in man performed in 2002). 5

20 2.1. Adverse events in medicine Historically, medical workers have been expected to function without error (de Leval 1997). As concluded in the 2004 Health Grades report, Patient Safety in American Hospitals, over 575,000 preventable deaths occurred as a direct result of the 2.5 million patient safety incidents that occurred in U.S. hospitals from 2000 through 2002 (HealthGrades 2004). This estimated average of 191,000 deaths per year is nearly double the 98,000 annual deaths cited in the 1999 Institute of Medicine report, To Err is Human: Building a Safer Health System (Kohn et al. 1999). A startling percentage of these adverse events are associated directly with a surgical procedure. The Harvard Study (Brennan et al. 1991) found that patients were unintentionally harmed by treatment in almost 4% of admissions in New York State. The study found that 47.7% of adverse events were associated with a surgical operation. In 70% of patients the resulting disability was slight or temporary; however in 7% it was permanent with a 14% mortality rate due in part at least to their treatment. Serious harm therefore came to approximately 1% of individuals admitted to hospital (Sarker and Vincent 2005). The Utah Colorado Medical Practice Study provided further data on intra-operative events (Brennan et al. 1991). The annual incidence rate of adverse events in patients who received an operation was 3.0%. Among all adverse surgical events, 54% were deemed preventable Human factors and adverse events in medicine During the past 2 decades, the medical field has become increasingly aware of the need to understand the human factors behind adverse events (Carthey et al. 2001). Of particular importance is research that takes a systems approach to identify the organizational, team, equipment, and human pressures that lead to adverse events (Ummenhofer et al. 2001). The majority of errors made in patient care are not the result of isolated individuals actions, but rather from the deficiencies in the systems of which they are a part of (Carayon et al. 2006, 6

21 Reason 2000). The top-level system approach is that humans are fallible and errors are to be expected, even in the best organisations. Errors are seen as consequences rather than causes, having their origins not so much inherent in human nature but in upstream systemic factors (Reason 2000) Adverse Events in EVAR Short term results of EVAR are favourable when compared to open surgical repair. EVAR is associated with a reduced short term mortality risk (1.6% EVAR, 4.6% open repair), however, the long term durability of stent grafts has been disputed (Doyle et al. 2008). The main drawback of EVAR is the higher proportion of re-intervention required compared to open-repair; 29.6% compared to 18.1% respectively at 6 year follow up (De Bruin et al. 2010). Endoleaks The most common adverse event by far are endoleaks, characterised by persistent blood flow within the aneurysm sac. Endoleaks are classified into 5 categories with type II the most common (retrograde flow into the aneurysm sac from lumbar and/or visceral arteries). Approximately 15-20% of EVARs are complicated by endoleaks with up to 12% requiring additional procedures to manage them (Faries et al. 2003). Up to 10% of patients require intervention to correct a type I endoleak (blood flow between the stent graft and aorta) detected on 30 day surveillance CTA (White and Stavropoulos 2009). The cause of type I endoleak is usually due to suboptimal positioning of the graft, poor seal between the neck of the aneurysm and graft, or distal migration. Type I endoleaks are in direct communication with the arterial circulation and require urgent intervention. If arising from migration or misplacement then an aortic cuff or iliac extender(s) can be used to extend the landing zone. Migration The risk of stent graft migration increases over time as pulsatile blood flow exerts a downwards force on the graft. This can result in exposure of the sac to 7

22 systemic pressure and increase the likelihood of rupture. There are multiple preoperative and postoperative causes of migration e.g. hostile anatomy (such as severe neck angulation) and short landing areas (particularly if the aneurysm involves the renal arteries). Medical device design can impact the rate of migration, due to inadequate fixation (active or passive) and short overlaps of contralateral legs and/or extenders/cuffs. Changes in the aneurysm morphology including sac expansion as a result of endoleaks or even significant reduction due to thrombus reabsorption can also alter the device position and orientation. Studies have shown that one of the main predictors of stent migration is the length of proximal fixation; each millimetre of increased fixation reduces the risk of migration by 2.5% (Zarins et al. 2003). Adequate overlap is required on joints between graft bodies to prevent type III endoleaks and is an additional reason why extreme precision is essential when deploying a stent graft (Ilyas et al. 2015) Adverse Events in TAVI Vascular complications, paravalvular leaks, stroke, and conduction disorders requiring Permanent Pacemaker Implantation (PPI) are the most prevalent complications of TAVI procedures (Van de Veire 2010). Appropriate patient selection with characterisation of vascular anatomy and aortic annulus size should be carefully evaluated. Exact valve placement is the most crucial step during the intervention, as other clinical complications can arise from a misplaced valve which are difficult to manage and require different bailout strategies. The following complications were reported from a systematic review of the safety and clinical effectiveness of TAVI by Yan et al. (2010): -Conduction disturbances (High-degree AV block (10-30%), -paravalvular leak (4-35%) -cardiac tamponade (1-9%) -coronary Ostia occlusion (0.5-1%) 8

23 -aortic dissection (0-4%) Malposition of the prosthesis rarely occurs (i.e. antegrade migration), however, there was a 5.3% incidence (9/170 patients) reported by Ali et al. (2008). Conduction Disturbances Conduction disturbances have widely been reported as the most prevalent complication during TAVI. A study by Baan Jr et al. (2010) reported the postprocedural incidence of new Left Branch Bundle Block (LBBB) of 65%, which is comparable with previous studies that reported a new LBBB incidence of 40% to 60% after TAVI (Calvi et al. 2009, Jilaihawi et al. 2009). All new LBBB in the cohort developed during the procedure, with most occurring directly after expansion of the prosthesis; therefore new-onset LBBB can be attributed solely to the expansion of the prosthesis imposing on the left bundle conduction tissues (Baan Jr et al. 2010). The only factor shown to be predictive for newonset LBBB is the depth of the prosthesis implantation in the LVOT: 10.2 ± 2.3 mm in the new-lbbb group versus 7.7 ± 3.1 mm in the non-lbbb group (P =.02). A summary of the main findings from the literature supporting implantation depth as a predictor of PPI is presented in Table 1 overleaf. The next section presents an overview of the human factors theoretical models used to investigate factors with the potential to cause adverse events during endovascular procedures. 9

24 TABLE 1: SUMMARY OF PERTINENT LITERATURE ON TAVI ADVERSE EVENTS Author No. of Patients Main Findings (Binder et al. 2013) 89 Low THV implantation was associated with new left bundle branch block and complete heart block (3.4 ± 2.0 mm vs. 5.5 ± 2.9 mm, p = 0.01) and with the need for permanent pacemaker implantation (3.5 ± 2.0 mm vs. 7.1 ± 2.5 mm, p = 0.001). In contrast, labelled THV size and THV area oversizing was not associated with atrioventricular conduction disturbances. (Muñoz- García et al. 2012) (Gilard et al. 2012) (Meredith et al. 2012) (Leon et al. 2011) 195 After the implantation, 48 patients (27.6%) required a definitive pacemaker due to AV conduction disturbances. The depth of the prosthesis in the LVOT was greater in the TS patients than in the Accutrak patients (9.6 +/- 3.2 mm vs /- 3 mm, p <0.001). Only 36.3% of the TS patients had a depth <8 mm compared with 74.6% of the Accutrak patients (p < 0.001) (Fig. 1). The need for a pacemaker was greater in the TS patients than in the Accutrak patients (35.1% vs. 14.3%, p= 0.003). New-onset LBBB occurred in 41.5% of patients with TS and 56.3% of patients with Accutrak system (p = 0.210) after excluding the patients with LBBB on baseline ECG. With the introduction of the new Accutrak release system, we noted a significant reduction in the need for a pacemaker after the percutaneous implant with the CoreValve aortic valve prosthesis, falling from 35.1% to 14.3%. The main reasons for this decline were the simplicity to position the CoreValve prosthesis higher and because of fewer manipulations and less trauma to the LVOT during the procedure with the Accutrak release system A new, permanent pacemaker was required more often in patients receiving CoreValve devices than in those receiving SAPIEN devices (24.2% vs. 11.5%).In our study, one of the most common complications was the need for a permanent pacemaker (in 15.6% of patients). In the other registries, the rates were similar, with 13% in Belgium, % in the United Kingdom, % in Italy,11 and 6.7% in SOURCE.13 In the PARTNER trial, the reported rates were 3.4% for cohort B18 and 3.8% for cohort A. The proportion of patients who needed a permanent pacemaker was highest in the German registry (39.3%). The higher rate of pacemaker placement with CoreValve implantation than with SAPIEN implantation in our study confirmed the findings in previous studies 428 PPM rates were evaluated by experience; for <30 implants, the PPM rate was 33.1%; for >30 implants; 19.9%. N/A Consensus Report In early experiences with TAVI, new-onset bundle branch block has occurred in up to 45% of patients and the need for permanent pacemakers has varied from as low as 4% to as high as 33% (95 98, ). Differences among devices and heterogeneity in physician and country-based health- care thresholds may explain the significant inter-hospital variability in new permanent pacemaker requirements after TAVI. 10

25 (Bosmans et al. 2011) (Moat et al. 2011) (Muñoz- García et al. 2010) (Baan Jr et al. 2010) (Petronio et al. 2010) (John et al. 2010) (Pasic et al. 2010) (Eltchaninoff et al. 2010) 328 New Permanent pacemaker 9/181 (5%) Edwards Sapien. 31/138 (22%) Corevalve. 877 The requirement for a new permanent pacemaker was significantly more common with CoreValve implants than with SAPIEN (24.4% vs. 7.4%; p < ). Some degree of paravalvular AR (angiographic grade 1) occurred in 61% of patients(moat et al. 2011), with this being moderate to severe (AR >2) in 13.6%. Moderate to severe leaks were significantly more common with the CoreValve device. 65 The need for a permanent pacemaker was associated with a greater depth of prosthetic implantation in the LVOT, with a mean distance of 13 (2.5) mm in the patients who required a permanent pacemaker as opposed to 8.8 (2.8) mm in those who did not require one (P<.001). No differences were found concerning the dimensions of the aortic annulus or the prosthesis/annulus ratio, the degree of left ventricular hypertrophy or valve calcification. The patients with RBBB required a pacemaker more often than those who had LBBB. Furthermore, the depth of the prosthesis in the LVOT was the only predictor of the need for a pacemaker in the multivariate analysis (OR=1.9; 95% CI, ; P=.007) 34 Percutaneous aortic valve implantation with the CoreValve prosthesis results in a high incidence of total atrioventricular block requiring PPI and new-onset LBBB. Pre-existing disturbance of cardiac conduction, a narrow left ventricular outflow tract, and the severity of mitral annular calcification predict the need for permanent pacing, whereas the only factor shown to be predictive for new-onset LBBB is the depth of prosthesis implantation 514 The most common in-hospital complications were the onset of a left bundle branch block (22.4%) and the need for pacemaker (16.3%). 100 All of these patients underwent post-dilation. In 4 of these patients a repositioning manoeuvre with a snare catheter was attempted, due to an additional so-called deep position, which caused suboptimal results after post-dilation. This snare manoeuvre improved the AR in 1 patient; the remaining 3 patients underwent a second implantation of a CoreValve prosthesis (valve-in-valve replacement or 2 valves in series ), which improved the regurgitation grade in all cases (Table 6). 175 Ten patients (5.7%) required pacemaker implantation because of higher-grade aortic valve block after surgery. 244 A pacemaker was implanted in 29 patients (11.8%), more frequently after CoreValve implantation (P = 0.001). First-degree atrioventricular block was present in 2 of the 29 patients (6.9%) needing a pacemaker vs. 20 of the 215 patients (9.3%) who did not need a pacemaker (P = 0.67). Left bundle branch block was present in 4 of the 29 patients (13.8%) needing a pacemaker vs. 23 of the 215 patients (10.7%) who did not need a pacemaker (P = 0.62). 11

26 (Avanzas et al. 2010) (Calvi et al. 2009) (Zajarias and Cribier 2009) (Webb and Lichtenstein 2008) 108 Definitive pacemaker implantation was carried out for atrioventricular block in 38 patients (35.2%) 30 Overall, 17 of 25 patients had at least one new onset CD at the time of discharge from the hospital, representing a cumulative incidence of 68%. Complete AVB is a complication of PAVR which might be related to the anatomic variability of the atrioventricular node. Specifically, being located on the right side of the interventricular septum, near the left coronary cusp, the atrioventricular node may be implicated when the prosthesis is expanded, though other clinical factors may also be involved. 355 Permanent pacemaker implantation was required in 9.3% and seems to be a function of prosthesis position N/A Review Paper Corevalve implantation is more frequently associated [with] early and late atrioventricular block, presumably due to the greater extension into the left ventricular outflow tract with compression of the septal tissues. The requirement for new pacemakers was 3-fold higher following Corevalve as compared with Sapien implantation in both the French and English national registries. 12

27 Human factors theoretical models Human factors is the scientific discipline concerned with the understanding of interactions among humans and other elements of a system and the application of theory, principles, data, and methods to optimise human wellbeing and overall system performance. Arguably the most widely known and applied models in the realm of patient safety for over 25 years is the Swiss cheese model. While Reasons model is widely referred to in the literature it offers little by way of guidance to researchers as a method of analysis identifying latent or active errors that could potentially trigger an incident or adverse event. The model itself has been described more as a conceptual analogy, imparting a basic understanding of how several minor errors can manifest into an adverse event within healthcare organisations (Figure 1). In the model there are several layers between the hazards and losses (in this case an adverse event). Each layer represents an additional barrier to prevent an error progressing. Under certain conditions however, a progression of missed opportunities can result in an error becoming an adverse event. The term error can be applied to any planned actions that fail to achieve their desired effect without the intervention of some chance agency (Reason et al. 2006, Reason 1990, Reason 2000). These errors can be categorised into two broad variants: - Slips & Lapses: Where an action does not go according to plan. - Mistakes & Adverse Events: Where the plan itself is inadequate to achieve its objectives. 13

28 FIGURE 1: SWISS CHEESE MODEL OF ACCIDENT [ADVERSE EVENT] CAUSATION (REASON 2000) The human information processing model (Figure 2), (Hollands and Wickens 1999) visually represents the allocation of our attention resources to carry out a task. External factors, especially in a surgical context (personal issues, hospital pressures, patient factors, theatre team etc.) can impact on attention resources leading to a lapse or mistake. These external pressures however, are often overlooked by the primary operators in the healthcare setting. A study by Sexton et al. (2000) to compare attitudes towards stress, error and teamwork in medicine and aviation. They found that 70% of surgical consultants felt that even when fatigued, they perform effectively during critical phases of operations/patient care and 76% feel that their decision making ability is as good in medical emergencies as in routine situations. By comparison, when pilots were asked the same questions the response rate was 26% and 64% respectively. Sarker and Vincent (2005) sought to further refine the combination of human error and performance models in a surgical specific model. Figure 3 outlines their proposed theoretical model of how errors may occur in surgery. External stressors, compounded by small slips and lapses can eventually lead to the most extreme form of adverse event, patient death. 14

29 These models provided the foundation for the development of the System Engineering Initiative for Patient Safety Model (SEIPS Model), which was selected as the theoretical human factors model for this research. The SEIPS model is discussed in detail in the next sections. FIGURE 2: HUMAN INFORMATION PROCESSING MODEL (HOLLANDS AND WICKENS 1999) FIGURE 3: MODEL OF FACTORS INFLUENCING SURGICAL ERRORS (SARKER AND VINCENT 2005) 15

30 Systems Engineering Initiative for Patient Safety (SEIPS) While patient safety research clearly recognises the need for human factors engineering and a systems approach in research, there are few models developed solely for the evaluation of health care systems. The SEIPS model was conceived as a response to a proposal by the Institute of Medicine (IOM) to apply human factors and systems engineering in order to improve the design of these systems (Carayon et al. 2006). The genesis of the SEIPS model lies in the work of Smith and Carayon-Sainfort (1989) on the balance theory of job design. According to balance theory, work organisation results in a system that has five core elements: -The individual -Task -Tools and Technologies -Physical environment -Organisation These five elements interact within the work system and can produce various physical and psychological stressors on the individual, producing a stress load. These stress loads detract from biological resources (energy expenditure, biomechanical strain), behavioural resources (motivation, coping behaviours) and psychological resources (perception, cognition, decision making, and emotion) (Hollands and Wickens 1999). The stress load, if maintained over a prolonged period of time can result in adverse outcomes for the individual, including compromised performance. Device design can directly impact the stress load an operator is subjected to, both positively and negatively. Well designed, easy to use and intuitive devices lessen the stress load while counterintuitive controls, confusing instructions/markings and uncomfortable user interface can add an avoidable burden to the operator. This is especially pertinent during a high pressure situation such as a rapid decline of a patient s 16

31 condition The physical responses to stress loads are directly influenced by the individuals physical capacity, health status and motivation. The psychological responses are influenced by past experience, personality and the social situation. According to balance theory, the effects of the work system on the individual are assumed to be mediated by the stress load (both physical and psychological). The SEIPS model is an expansion of, and improvement on, the original balance theory model and Wickens model of human information processing. In order to characterise the interactions between people and their environment in a concise manner the SEIPS model uses and builds on the nomenclature used in the balance model developed by Carayon and Smith (2000). The SEIPS model is summarised as: An individual performs a range of tasks, using various tools and technologies. The performance of these tasks occurs within a certain physical environment and under specific organisational conditions SEIPS 2.0 In recognition that the discipline of human factors and the healthcare domain are evolving, SEIPS 2.0 was developed by Holden et al. (2013). The expansion of SEIPS included an additional core element to the model. External environment is included which encompasses macro-level societal, economic and policy factors outside of an organisation (i.e. the insurance or governmental scheme available to an individual that can influence the level of care they receive). The components of the SEIPS 2.0 model are: -Individual -Task -Tools and Technologies -Internal Environment -Organisation 17

32 -External Environment Therefore the SEIPS 2.0 model can be expressed as: An individual performs a range of tasks, using various tools and technologies. The performance of these tasks occurs in a physical environment, under specific organisational conditions within the wider socio-economic environment. SEIPS 2.0 introduces further developments on the original, as discussed here: Configuration A further addition to the original SEIPS model is the inclusion of configuration. The concept underpinning configuration is that while all of the components of the healthcare system have the potential to influence each other, only a specific subset of influencing factors may be active together in any given situation. In the SEIPS 2.0 model, any number of work system components can interact simultaneously at a moment in time to shape performance, processes and outcomes Engagement Engagement in the SEIPS 2.0 model is a method to decompose work activities based on the individuals actively engaged in performing them. To be engaged is referred to as being an active agent who performs some or all of the health related work activity. Indirect or passive contributors to the work activities are referred to as co-agents. It is important to note that these agents are assigned at a given moment in time; therefore many individuals may be assigned as agents for a given process including healthcare professionals, patients or family members etc. 18

33 Professional work Professional work is a configuration whereby those engaged as primary agents are professionals (or team of) with minimal patient, family or non-professional involvement. A brief example is described in Table 2 overleaf for the transcatheter replacement of an aortic valve on a sedated patient. In the example, the agents are the members of the surgical team working together. The patient and family are deemed to be co-agents as they are not able to participate in the surgery, even if they may have had active engagements earlier, for example, in deciding, communicating and/or preparing for the surgery. This break down of the work process provides a snapshot that captures a moment in time; the TAVI procedure itself. Using SEIPS 2.0 as a conceptual framework, it is possible to define a configuration of work system components in order to evaluate and improve specific elements effecting performance, processes and ultimately, outcomes for the patient. The next section details the SEIPS model when applied to endovascular interventions, and identifies the most pertinent area for exploration in this work. 19

34 Table 2: SEIPS applied to TAVI, Adapted from Holden et al. (2013) Professional Work Example Agent(s) Co-Agent(s) Work Systems Factors Process Outcomes Work in which a healthcare professional or team of professionals are the primary agents, with minimal active involvement of patients, family caregivers and other non-professionals. Cath Lab team performing TAVI Cardiologist, Radiographer, Cardiac technician, ~3Staff Nurses. Patient, Relatives Person(s) factors include skill levels of all the involved parties, experience with the procedure and professionals personal preference concerning the procedure (e.g. preferences for tools and supplies, use of time outs and patient transfer processes). Task factors include the difficulty of the case and the familiarity of work tasks for various team members. Tool/technology factors include the availability or usability of patient monitoring technologies and patient checklists, the delivery system for the TAVI valve, and various other medical devices (i.e. intra-aortic balloon pump on standby). Organisation factors include the number of hours or surgeries worked per day by the team members, whether work-arounds need to be used due to lack of personnel, whether all team members can work in unison and can speak up and the availability of appropriate detailed procedures for emergency situations. Internal environment factors include operating room hygiene, lighting, air quality, noise, workspace design and layout, and operating room size. External environment factors may be the impact of budget and cost on the quality of the tools/technologies used, marketinfluenced pay levels for personnel, and societal expectations for patient and family preferences. These factors interact to shape surgical performance The process of endovascular replacement of an aortic valve of a patient includes applying the anaesthetics to the patient, surgical skin preparation to prevent infections, inserting central line(s), opening the femoral artery, crossing the aortic arch, ensuring correct positioning, delivering the aortic valve, placing pacing wires and closing the patient. Performance of each may be shaped by unique configurations of work system factors. Proximal outcomes include successful completion of the surgery, minimal errors and adverse events (such as permanent pacemaker implantation), and surgical team member stress and fatigue. Distal outcomes include full recovery of the patient, patient satisfaction with their care and trust in the healthcare delivery system, no downstream complications (e.g. healthcare-associated infections), job satisfaction of surgical team members and long-term profits for the institution 20

35 Use of the SEIPS model in this research As discussed previously, a subset of all elements in the SEIPS model combine at any one moment to shape the clinical outcome for the patient. For the purposes of this research, it is useful to evaluate each element of the SEIPS model in the context of an endovascular procedure to identify any potential areas to improve patient safety: -Individual The individual (operator) in this instance is primarily a consultant (cardiologist, Interventional Radiologist, Vascular Surgeon etc.). Consultants are highly experienced and specialised with a wealth of tacit knowledge and are experts in their respective fields. -Task The task relates to the level of difficulty, complexity, variety, and ambiguity. In the context of endovascular intervention, task relates to the anatomical variability from patient to patient. These differences add significant difficulty to each case. -Tools and Technologies Tools and technologies in healthcare describes a wide range of devices and equipment. Adhesive plasters to high end imaging systems and patient record systems all fall under the umbrella of this element. Specifically in endovascular interventions, the primary tools are the delivery systems used in the specific intervention. -Internal Environment The internal environment is primarily concerned with the immediate vicinity, in the Cath lab or operating room, cleanliness, lighting, air quality, temperature, noise and layout are typically well controlled in the hospital environment. -Organisation Organisational factors are those structures that are broadly external to the individual that organise work schedules, training, policies and resource 21

36 allocation. In the healthcare setting organisational mainly refers to management and administration elements rather than clinical considerations. -External Environment External environment relates to the macro level socio-economic stresses that either directly or indirectly effect the delivery of care to a patient. These include government spending policy for public hospitals, private insurance cover levels and wider global considerations such as recession. Within the SEIPS model, improvements affecting patient safety are possible in all elements. However, some require significant modifications in attitude, policy and macro-management, and as such they fall beyond the scope of the current research. The individual in the SEIPS model is a highly skilled, driven and motivated person, often with decades of intense training. The task cannot be altered or modified as it is a function of anatomical diversity. It is possibly the least applicable element in the context of endovascular intervention. The internal environment is carefully controlled, with only severe breaches of protocol affecting the status quo. The organisation is a deeply engrained sociotechnical system, formed over a prolonged period of time and subject to little change due to adherence to fixed rules and a hierarchy of authority. The external environment is at such a macro level, only governmental or global movements can effect change. Therefore, the most opportune area to improve patient safety during endovascular intervention is Tools and Technologies. The design and function of endovascular devices directly impacts the likelihood of an adverse event occurring. In the next section, the role of design in facilitating/negating adverse events is discussed. 22

37 2.2. Role of medical device design in preventing errors Technologies often fail to deliver their promised benefits when they are not designed in a way that matches the needs, cognitive processes, and environments of the intended users (Fairbanks and Wears 2008). In addition, purchasers (e.g., hospital supply officials) and end users do not realise the importance that device design can play in enhancing or degrading safe and effective performance (Johnson et al. 2007). If one takes a systems approach to designing a medical device in a user centred manner, then the device should match the way practitioners think, the way they operate in their daily practice, and the limitations of their environment and working conditions (Klien et al. 2004). Although the FDA requires a human factors analysis during the approval process for new medical devices, these analyses are neither independent nor publicly available (Burlington 1996). High quality, well-designed medical devices are critical to ensuring safe and effective treatment for patients as well as to ensure the safety of all end users. Medical devices are a wide ranging group of products ranging from the most simplistic of adhesive plasters to incredibly complex pieces of equipment such as Cardio-Pulmonary Bypass (CPB) machines (Martin et al. 2008). The ultimate goal of any medical device, regardless of complexity, is to improve the wellbeing of person receiving treatment (Sharples et al. 2012). The application of human factors methodologies has significantly improved the safety of medical devices, particularly infusion pumps. The success of this research has led to regulatory bodies such as the FDA and EC to require manufacturers to apply human factors methods as part of the design process (Martin et al. 2012). The design process for medical devices is of course, heavily regulated by the FDA for all products that are to be sold in the US. While the regulatory requirements are notably different in the European and American markets (particularly the 23

38 requirement to show efficacy in the US), it is generally accepted that following the FDA regulations makes securing CE making almost certain Medical device design controls Design controls are a system of checks and evaluation that are incorporated into the design process of medical devices. The application of design controls is the systematic assessment of the design as a fundamental part of the design process (FDA 1997). In this way, discrepancies between the proposed design solution and the requirements of the device are identified, and remedied, early in the design process. The application of design controls increases the potential for a device to make it to market and to fulfil its intended use effectively. Device controls are one component of a comprehensive quality system that extends to the entire life of the product, from disposable syringes to complex robotic-assisted surgical systems (FDA 1997, FDA 2011) Application of design controls The design of endovascular devices, as all medical devices, is controlled in the US by the Code of Federal Regulations, specifically: Title 21, Chapter I, Subchapter H, part Quality System Regulation (21 CFR 820). Part , Design Controls sets out the main components of the process for documenting the development of a medical device, (the unabridged text of which can be found in Appendix IV) section (a) (General) states: (1) Each manufacturer of any class III or class II device, and the class I devices listed in paragraph (a) (2) of this section, shall establish and maintain procedures to control the design of the device in order to ensure that specified design requirements are met. Design controls may be applied to any new product development process however they are of particular importance in medical device design. Figure 4 shows a high level overview of how design controls are applied throughout the 24

39 medical device development process. It also illustrates the cyclical nature of the process. User needs are used to define design inputs, this is accomplished through the use of formative testing (discussed in detail in section 2.3.1). These inputs are used as a benchmark to test design outputs (design solutions) against. This testing process is repeated iteratively until an optimal design solution is arrived at. The finished medical device is then tested against the user needs through summative testing (discussed in section 2.3.2). All of this testing schedule must be documented in the design history file, which is the required document to fulfil 21 CFR Figure 5 graphically illustrates the complex relationship between 21 CFR (design controls), ISO (risk management), and IEC (Usability engineering) standards. Each of the standards, with the addition of several field specific standards (e.g. ISO medical electrical equipment, ISO 7197 neurosurgical implants) must be completed as partial fulfilment of the other. FIGURE 4: APPLICATION OF DESIGN CONTROLS TO WATERFALL DESIGN PROCESS (MEDICAL DEVICES BUREAU, HEALTH CANADA) 25

40 FIGURE 5: RELATIONSHIP BETWEEN RISK MANAGEMENT AND USABILITY ENGINEERING STANDARDS 26

41 User centred design of medical devices User Centred Design (UCD) is a term to describe a design process whereby the end user influences how the design evolves and meets their needs. The term UCD originated in Donald Normans research laboratory at the university of California San Diego in the 1980s (Abras et al. 2004). Norman built further on the concept of UCD in his seminal book: The Psychology Of Everyday Things (POET). In POET, Norman emphasizes the needs of the user, and the usability of the design. He suggests four basic heuristics to influence design: Make it easy to determine what actions are possible at any moment. Make things visible, including the conceptual model of the system, the alternative actions, and the results of actions. Make it easy to evaluate the current state of the system. Follow natural mappings between intentions and the required actions; between actions and the resulting effect; and between the information that is visible and the interpretation of the system state (Norman 1988). These heuristics place the end user at the centre of the design process. The role of the designer in this process is to ensure that the user is able to use the product to carry out the task with effectiveness, efficiency and satisfaction (Abras et al. 2004). The benefits of UCD are vast and encompass the identification of new ideas, design directions, better user experience, reducing complaints/error, and a more precise definition of functionality (McClelland and Suri 2005). Hallbeck (2010) Portrays UCD as both a philosophy and a process and in a review of approaches and standards notes that most medical devices, including surgical and laparoscopic tools, have not been designed using UCD principles; in fact some appear not to have considered there was a user. 27

42 Major advantages of utilising the UCD approach is a deeper and more robust understanding of the psychological, organisational, social, and ergonomic factors that emerge from the involvement of users from the very beginning of, and throughout the product development process. The involvement of real life end-users ensures that the product will fulfil its intended purpose in the scenarios and environment in which it will be used. The UCD approach leads to the development of products that are more effective, efficient and safer (Abras et al. 2004). ISO defines user centred design as: - An approach to design that is characterised by the active involvement of users, a clear understanding of user and task requirements, an appropriate allocation of function between users and technology, iterations of design solutions, and multi-disciplinary design. Abras et al. (2004) discuss some of the methods that have subsequently been developed to support UCD such as usability testing, usability engineering and participatory design. The standard goes on to address the common held misconception that usability and user centred design are one and the same usability engineering is often used as a substitute for user centred design. However applying usability engineering methods does not necessarily prescribe the active user involvement that is the essence of user centred design. In addition, usability engineering often over-emphasises the role of evaluation methods. Human-centred design, on the other hand, refers to the process of analysing context of use, eliciting user requirements, producing design solutions and evaluating the design against the requirements, all in an iterative fashion (ISO :2010 Definition 4.8). It can be said that usability is a measure of success in implementing user centred design. 28

43 Usability The term usability was coined in the early 80s in order to replace the term userfriendly which had acquired a host of vague and subjective connotations (Bevana et al. 1991). Usability can be described as the development of devices (products) which are fit for use (Matern and Büchel 2011). Originally envisioned as a tool to aid and evaluate the design of human-computer interfaces, the application of usability has broadened significantly over the last three decades. In recent years the focus of design has moved from the reduction of physical exertion to the reduction of mental exertion while using a product. It is also important that the product (device) should not distract the users attention away from the actual task as a result of the design. Devices which are intuitive, easy to use and satisfying are described as usable. Usability is a qualitative attribute that indicates how easy a product is to use (Matern and Büchel 2011). Telling designers that a device should be intuitive is not enough, therefore design principles are necessary to guide them (Abras et al. 2004). Norman (1988) Put forward seven principles of design [ ] essential for facilitating the designer s task. Normans work concentrated on the necessity to explore needs and desires of the intended users fully. The need to involve actual users in real world scenarios and environments during every stage of the design process was a natural evolution of the concept of user centred design. Nielsen (1993) Adapted and popularised Normans design principles by stressing the importance of considering usability as a multifaceted approach, describing five usability attributes: Learnability, efficiency, memorability, errors, and satisfaction. The application of HF to the design of medical devices strives to deliver robust, safe and effective products to the market. To ensure this, usability testing is prescribed in the relevant standards throughout the design process of medical devices, and is discussed in the following sections. 29

44 2.3. Usability testing of medical devices during design Medical device users are an extremely heterogeneous group. They range from exclusively physicians, to nursing staff, support staff, patients, to their family and carers. Measuring and fulfilling user requirements during medical device development results in successful products that improve patient safety, improve device effectiveness and reduce product recalls and modifications (Jennifer et al. 2006). Usability testing should not be considered as product validation. Usability testing can verify that the design outputs and design inputs correspond; that is that the design solution encompasses all of the design specifications stated at the beginning of the project. Medical device validation should include laboratory testing to evaluate device efficacy, reliability, safety and performance before clinical testing (Weinger et al. 2010). The need for usability testing is mandated in the US by the FDA in 21 CFR 820 Section Regulatory Basis for HFE/UE Analysis and Testing: Human factors techniques play an important role in fulfilling the design control requirements of the Quality System regulation, 21 CFR Part 820. Specifically, human factors testing helps ensure proper design of the user interface [delivery system]. The risk analysis that fulfils quality system requirements should include use error. To establish the design input for the user interface and carry out design verification, human factors activities conducted throughout the development process can include task/function analyses, user studies, prototype tests and mock-up reviews. Formative and validation testing fulfil the requirements to test the device under realistic conditions. Validation testing should be used to demonstrate that the potential for use error has been minimized. While usability testing is deemed essential to the development of robust and safe medical devices, there are no formal guidelines on how to integrate this testing into the design process. There is a complicated relationship between the traditional design process and the regulatory requirements for safety, documentation and quality (as discussed in section 2.2.1). The author collaborated with Design Partners, (Bray, Wicklow, Ireland) to create a framework to implement the required usability testing within their medical 30

45 device design process. The model (presented as a fold out overleaf) was presented at the Human Factors and Ergonomics Society World Congress 2014 in Chicago (O'Sullivan et al. 2014). 31

46 FIGURE 6: INTEGRATING HUMAN FACTORS INTO THE DESIGN PROCESS 32

47 Formative testing Formative usability testing is performed early in the design phase using simulations and early working prototypes. Formative testing is used to guide medical device design development and elicit information via user interaction with said device. These formative tests can be relatively informal, particularly during the early stages of design and can identify problems that were not observed or fully understood using analytical methods (contextual inquiry, focus groups, task analysis interviews, and heuristic review). Formative testing can be applied to different Exploratory testing Tests users performing high level or walkthroughs of tasks using low fidelity simulations (e.g. paper or foam mock ups) Assessment testing Tests that provide users with more realistic tasks to perform using more developed devices and/or simulations (e.g. testing the feel and control of a device) Comparison (contrast) testing Tests comparing two or more design solutions (e.g. testing clockwise and counter clockwise actuated controls for TAVI systems) Comparison (competitor) testing Tests that gather performance characteristics of competitor products. These tests can form part of the fuzzy front end of the design phase to distil the most desirable features of predicate devices. Small numbers of test participants and the facility to test multiple iterations of solutions are paramount to garnering the most benefit from formative testing. This rapid iterative approach helps to inform design decisions, training requirements and the content of the Instructions For Use (IFU). FDA guidance emphasises the use of formative testing according to Table 3. 33

48 TABLE 3: HFE GOALS OF FORMATIVE TESTING (FDA 2011) Formative testing -Identify and prioritize tasks according to relative risk to the user beyond estimates derived from analytical techniques; -Guide development of use scenarios to be employed during subsequent design validation testing; -Identify use-related hazardous situations leading to the development of risk mitigation strategies; -Evaluate trade-off considerations and effectiveness of design enhancements, training and instructions for use; -Guide modification of the device design to optimize the user interface with respect to device safety and effectiveness; and -Clarify the dynamics of device-interaction associated with known or suspected use error scenarios Summative (validation) testing Summative testing is performed in the latter stages of the design process as part of device verification and validation. Summative testing requires formal acceptance criteria according to the design inputs established at the beginning of the design phase (see Figure 4 in section 2.2) (Weinger et al. 2010). Validation testing is used to establish safe and effective performance of a medical device using actual users in realistic conditions. The use of iterative formative usability tests, performed throughout the early and mid-stages of the design process should ensure that there are little or no usability issues discovered during summative testing. Simulated use testing is identified by both the IEC and FDA standards as an appropriate tool for formative and validation testing of medical devices: FDA Draft Guidance on Applying Human Factors to Medical Devices The FDA guidance document, as controlled under 21 CFR (section 2.2.1) lays out the requirements for validating medical devices, and makes reference to simulated use in: Section 10 Human Factors Validation Testing 34

49 The human factors validation test demonstrates that the intended users of a medical device can safely and effectively perform critical tasks for the intended uses in the expected use environments. It is particularly important during validation testing to use a production version of the device, representative device users, actual use or simulated use in an environment of appropriate realism, and to address all aspects of intended use. Simulated use is discussed as a separate section in the standard: Section 10.1 Simulated Use Validation Testing The conditions under which simulated use testing is conducted should be sufficiently realistic to enable the results of the testing to be generalized to anticipate actual use. The need for realism is therefore driven by the analysis of users, use environments, the device user interface and intended uses. To the extent that environmental factors are found to affect user performance, they should be incorporated into the simulated use environment (e.g., dim lighting, multiple alarm conditions, distractions, multi-tasking and workload). Application of Usability Engineering to Medical Devices (IEC 62366) IEC advocates the use of simulation to explore worst case scenarios, complex failures and environmental interactions while validating medical devices specifically section: D.5.13 Simulated clinical environments and fieldtesting Simulated clinical environments permit evaluation in a controlled manner in a setting containing some or all of the essential attributes of the actual clinical environment for which the MEDICAL DEVICE is being designed. Simulations facilitate creation of worst-case USE SCENARIOS and complex failures. A high- RISK MEDICAL DEVICE or one involving tasks that are more complex can be tested in high-fidelity simulators, such as a full-scale, simulated operating room with functional manikin. High-fidelity simulation allows the test team to evaluate dynamic interactions among multiple MEDICAL DEVICES, personnel, and task constraints Simulation in usability testing of medical devices Simulation and simulated use testing are specified in US and European regulations as in appropriate human factors test method during medical device design. Simulation has long been adopted as an effective method of teaching 35

50 techniques and evaluation in many fields. Examples of professions that embrace simulation-based training include aerospace (flight simulators), the military (realistic war games), and nuclear engineering (systems simulators) (Halamek et al. 2000). Simulators are designed to recreate some aspect of the environment and / or the task at hand. In healthcare, this can range from silicone vein blocks for the practicing of venous puncture, through increasing levels of complexity and realism to the simulation of entire operating theatres (Maran and Glavin 2003). The term simulator when used in the context of healthcare usually applies to a device that s simulates a patient, or a part of a patients anatomy, and interacts with the actions of the user in an accurate manner without invasive procedures required (Gaba 2004). The development of human patient vascular simulators began in the late 1960s, accelerating in the late 1980s and early 1990s. Simulators are used to teach basic skills, such as respiratory physiology and cardiovascular haemodynamics, and advanced clinical skills, e.g. management of difficult airways. Simulation offers distinct educational advantages, especially for learning how to recognise and to treat rare, complex, unseen clinical problems. The costs of simulator-based educational programmes include facility, equipment and personnel (Good 2003). Central to the idea is usability testing for medical devices is the need to approximate real life conditions as closely as possible (as set out in both the FDA and IEC standards). For the treatment of aortic aneurysms and aortic stenosis, it is essential to evaluate both the prostheses and delivery system designs in a simulated environment before in-vivo testing (Sulaiman et al. 2008). In vivo animal models are used extensively as a means of evaluating new endovascular devices. These tests are, however, expensive procedures that usually only facilitate the use of one device per animal before it is euthanised. 36

51 Animal models rarely mimic the disease state that is being targeted. For example with aortic stenosis, prosthetic valves are implanted into aortic valves where no calcifications may be present, and with aortic aneurysms stent grafts may be deployed in the aorta but without the presence of an actual aneurysm. In both cases, evaluation of devices in a simulated anatomical model would provide usability testing that more closely resembles real world environments Suspension of disbelief The key to effective simulation based evaluation or training is achieving a suspension of disbelief on the part of the operator (Halamek et al. 2000). That is, there must be a sufficient level of fidelity for the operator to think and feel like they are functioning within a real environment. Simulators are controlled environments in which varying clinical scenarios can be experienced on demand. These scenarios can be scaled to fit the level of the operator, from engineers evaluating a prototype device to a consultant training on a new interventional technique. Some of the key advantages of simulators are shown in Table 4. Of critical importance to any simulator is the consideration of fidelity. Farmer et al. (1999) stated that fidelity is the extent to which the appearance and behaviour of the simulator/simulation match that of the simulated system, however confusion abounds over the ambiguous use of fidelity. TABLE 4: ADVANTAGES OF SIMULATORS (MARAN AND GLAVIN 2003) Risks to patients and learners avoided Undesired interference is reduced Tasks/scenarios cam be created on demand Skills can be practiced repeatedly Training can be tailored to individuals Retention and accuracy is increased Transfer of training from classroom to real situation is enhanced Standards against which to evaluate student performance and diagnosis educational needs are enhanced 37

52 Fidelity in simulation A recent attempt by Hamstra et al. (2014) to quantify the importance of simulator fidelity using a meta-analysis of over ten thousand articles was inconclusive. Fidelity appears to be subjective depending on what features of a simulator are emphasised or ignored, what the task is, or what the outcomes and objectives are. An example provided is that anaesthetists consider simulator mannequins as high fidelity for tasks such as intubation, whilst surgeons typically view cadavers or animal models of a superior fidelity to mannequins (Ocel et al. 2006, Yang et al. 2010). The disagreement about the concept of fidelity is not a new phenomenon or specific to health care. Beaubien and Baker (2004) note that in the early days of the field of industrial and organisational psychology fidelity was regarded as a binary concept i.e. consisting of high and low. The contemporary view is that this perspective is overly simplistic particularly in the sense that it emphasises the technology over the training goals, content, and design. A more useful framework for quantifying fidelity is to consider the work of Allen et al. (1991) who identified the two dimensions that underpinned a large variety of definitions of fidelity. Miller (1954) was the first to make the distinction between Engineering Fidelity and Psychological Fidelity in relation to simulation. These two definitions have been updated by Allen et al. (1991) and Hamstra et al. (2014) to that of Structural Fidelity (how the simulator appears) and Functional Fidelity (what the simulator does). Structural fidelity The degree to which the simulator or environment replicates the physical characteristics of the real task. Increasing the engineering fidelity invariably increases the cost of the simulator, and beyond a certain point, increasing the fidelity produces only marginal gains in performance over a simpler solution. For surgical procedures and medical device testing, structural fidelity is of high 38

53 importance for the task (e.g. when interacting directly with tissue or anatomies), while decision making and communication don t require the same degree of structural fidelity. Salas et al. (1998) concluded that the level of fidelity built into a simulator should be determined by the level needed to support learning of the tasks. Functional Fidelity The degree to which the skill or skills in the real task are captured in the simulated task. The level of fidelity required is dependent on the on the type of task and the level of training of the user. Highly complex simulators are inappropriate when training novices in basic skills, however, if developing fine motor skills, such as manipulating a guidewire trough tortuous anatomy, the simulator should be capable of reproducing the anatomical movements of the vessels. At advanced levels of training for complex tasks, simulators should provide real time feedback to support high level decision making (e.g. alarms). In the context of medical education, a low structural fidelity simulator with high functional fidelity may lead to a higher knowledge transfer (Boreham 1985). Practical differences between structural and functional fidelity An example of the ambiguity surrounding fidelity can be demonstrated by considering static animal tissue e.g., a fresh pig s foot. A pig s foot may be considered to be of low structural fidelity compared to a human hand until it is used to teach suturing techniques. At this point the tissue responsiveness offers precisely the level of functional fidelity required for the task of basic suturing skills (Hamstra et al. 2014). Furthermore if the pig s foot is draped in such a way that only a small portion of the epidermis is showing, there is a sufficient suspension of disbelief to facilitate the transfer of knowledge from the task. To that end it has been proposed that the terms structural and functional fidelity be replaced with Physical Resemblance and Functional Task Alignment respectively. The subtle shift in terminology acts to reinforce the 39

54 importance of the task and the need for active and intentional determination of the required task alignment Simulators used in medical contexts Simulators can range from the most basic cardboard representations for teaching task specific psychomotor skills, to high end virtual reality units capable of simulating various devices, anatomies and disease states. As structural and functional fidelity increase, so too does the cost of the simulator (Figure 7). For evaluation of endovascular devices and delivery systems, the predominant method in the literature involves using silicone anatomical models in pulsatile flow rigs (Doyle et al. 2008, Ene et al. 2011, Lynch et al. 2013, Poepping et al. 2004, Chong et al. 1998). These pulsatile simulators (e.g. Figure 8) while exhibiting high functional fidelity in their ability to reproduce excellent physiological conditions, do not provide high structural fidelity for usability testing of medical devices. The reason for these shortcomings are discussed in the next section. FIGURE 7: DIMENSIONS OF FIDELITY AND COST IN SIMULATION 40

55 FIGURE 8: TOP, AAA SIMULATOR (CHONG ET AL. 1998). BOTTOM, PULSATILE FLOW LOOP BY CABER, UNIVERSITY OF LIMERICK Silicone anatomical models Hollow silicone anatomical models are widely used for physical bench testing of medical devices, experimental models, and demonstration purposes. Silicones (such as Wacker RT601), have been shown to be good arterial analogous due to their non-linear behaviour at high strain loads (Doyle et al. 41

56 2008). Silicone models are also desirable for visualisation of devices inside the anatomy without the need for x-ray imaging as they are transparent. The lack of functional fidelity in silicone phantom anatomies has the potential to negatively impact the results of both formative and summative device testing. Hollow silicone anatomical models do not contain calcifications or thrombus, both of which have been shown to impact the response of anatomy and medical devices. The study of the importance of including both Intra-Luminal Thrombus (ILT) and calcifications in computer simulated models is well described in literature. Wang et al. (2002) demonstrated that the presence of ILT significantly impacts the distribution of peak wall stresses in AAA using finite element analysis of patient specific models. ILT has also been shown to increase the rate of growth of the aneurysm sack (Speelman et al. 2010). Li et al. (2008) confirmed the effects of ILT on wall stress and showed that the presence of calcifications increases peak wall stress, suggesting that calcification decreases the biomechanical stability of AAA. Traditional silicone models are limited to one material used to create the aortic wall. Thrombus may be added, but requires a third aluminium mold to be machined in order to produce a wax model of the true lumen first, as discussed below Process of forming silicone anatomical models Lost material casting is the primary method of manufacturing silicone vascular models. There are small variances, but the general method has been described at length in the literature (Poepping et al. 2002, Poepping et al. 2004, Barry J Doyle et al. 2008, Smith et al. 1999): A CNC milling machine is used to fabricate two concentric molds of the desired vascular geometry. The first is an aluminium mold (which represents the lumen boundary) milled with life-sized model geometry. A second mold (either aluminium or acrylic) is milled with a dilated model geometry that uniformly extends the life-sized geometry radially 42

57 by a specified amount to produce the desired wall thickness, in the case of aortic models, ~2 mm (Doyle et al. 2008), at all points offset to the inner lumen of each vessel. A lumen model is cast in the aluminium mold (sacrificial core), using either wax or a low-melting-point (47.2 C) alloy (such as: Cerrolow 117, Cerro Metal Products, Bellafonte, PA). The wax or metal cast is centred within the dilated-geometry mold, and silicone mixture is injected uniformly to fill the 2mm gap between the dilated mold and the sacrificial core. Once the silicone is cured, the dilated mold is removed and the sacrificial core is melted out, leaving a silicone model of the anatomy. Additional features such as plaques, narrowing s or ILT can be added by the use of subsequent molds and additional sacrificial cores. For AAA, these cores are inserted into the silicone aortic wall model and a softer grade of silicone injected into the aneurysm sack to replicate ILT (Figure 9). There is no mechanism to add calcifications to the aneurysm models produced by this method. This process can take from several days to weeks to complete a single anatomical model depending on the complexity of the model and the number of cores required. FIGURE 9: ADDITION OF ILT TO AAA MODELS (ENE ET AL. 2011) 43

58 Commercial silicone models Commercial Silicone arterial models are available from several suppliers. These models are either idealised or generalised from population data and do not include ILT or calcifications. The quotation below is taken from the same web page as Figure 10 and illustrates the dogma surrounding hollow phantom anatomies for device testing and training. Elastrat in-vitro models respect human anatomy and are designed for the education, development of stents, coils and catheters. They provide a realistic environment for the simulation of endovascular procedures, pre-surgery trainings, studies and teaching purposes for interventionists FIGURE 10: SILICONE AAA MODEL (A-S-A005+) BY ELASTRAT, GENEVA, SWITZERLAND Limitations of hollow silicone anatomical models The limitations of silicone models for use in usability testing of endovascular devices is the lack of functional fidelity. While using idealised models is appropriate for initial device design in the R&D laboratory, functional evaluation of both the device and delivery system is hampered. The main limitations of hollow silicone models for use during TAVI and EVAR evaluations are show in TABLE 5. 44

59 TABLE 5: LIMITATIONS OF HOLLOW SILICONE PHANTOMS FOR DEVICE TESTING Non tortuous anatomy No calcifications (PAD) No Calcifications - TAVI No Calcifications - AAA No ILT - AAA Vascular access via the femoral arteries is the basis of all endovascular interventions. Frequently current models are relatively straight forward Device performance, vascular access/complications, and procedural success can be influenced by heavy calcification/pad. Without the inclusion of calcification in the aortic valve, device deployment may show better outcomes than Calcifications have been extensively shown to effect both peak wall stresses and dynamic movement of AAAs by computational modelling Stent graft evaluation invalid due to floating nature of graft in an empty anatomical model. Neck angulation and landing zones are non-representative Imaging in endovascular simulation Endovascular procedures are carried out under fluoroscopy in the cath lab or hybrid operating theatre. At the formative stage of testing it is often impractical or prohibitively expensive to use a C-arm while evaluating early stage designs. The use of standard video equipment as an analogue of fluoroscopy in simulation is well discussed in the literature (Chong et al. 1998, Corbett et al. 2011), this method is inexpensive and effective when utilised correctly. Chong et al. used a plain film x-ray behind a silicone AAA model to recreate the effect of fluoroscopy while testing stent graft placement and migration. Use of simulated imaging increases both structural and functional fidelity of a vascular simulator. In clinical practice, an operator using an endovascular device will have their view focussed on an imaging screen above the patient. This is an important element of any endovascular simulator as direct visualisation of the device to be tested in an anatomical model diminishes the suspension of disbelief that facilitates effective simulation. Figure 11 shows the fluoroscopic view of an Edwards Sapien balloon expandable aortic valve being implanted. 45

60 FIGURE 11: FLUOROSCOPY IMAGE SHOWING A SAPIEN AORTIC VALVE BEFORE AND DURING BALLOON INFLATION (CHEUNG AND LICHTENSTEIN 2012) 46

61 2.5. Development of usability simulators for AAA and TAVI As discussed at the beginning of the literature review, two disease states were chosen as a focus for this research. EVAR the endovascular treatment of abdominal aortic aneurysms, and TAVI, the endovascular treatment of aortic stenosis. It is important to have a firm grounding in the disease states, causes, treatments, and adverse events in order to understand the importance of endovascular treatment options. A brief overview of both AAA and AS are presented below Abdominal Aortic Aneurysm (AAA) An aneurysm is defined as the localised and permanent swelling of a portion of a blood vessel (Figure 12) (Sakalihasan et al. 2005). While an aneurysm can theoretically form in any blood vessel, aortic and cerebral aneurysms are the most associated with the highest rates of mortality and morbidity. Aortic aneurysms are the most prevalent type of arterial aneurysm events, comprising ~85% of cases. Of these, Abdominal Aortic Aneurysms (AAA) are the most common. AAA affects between 5% and 9% of the population over the age of 60. It is responsible for over 10,000 death per year in the United Kingdom (~0.9% of all deaths) and is the 14 th leading cause of death in the United States FIGURE 12: ILLUSTRATION OF INFRARENAL AAA IN THE BODY (SOCIETY OF INTERVENTIONAL RADIOLOGY, 2010) 47

62 (Corbett et al. 2008, Kavanagh 2008, Figueroa and Zarins 2011). An AAA is typically defined as an aneurysm of the infra-renal aorta, however the aneurysm neck can commonly be supra-renal and involve the renal arteries themselves. While there is no definitive consensus on the definition of an aneurysm, a diameter greater than 30mm is generally accepted as a positive diagnosis. Other criteria have been proposed since the early nineties such as a diameter 1.5 times the expected diameter (Johnston et al. 1991). The incidence of AAA has increased rapidly over the past two decades, due in part to a number of factors; ageing population, smoking, screening programmes and improved diagnostic tools. Most AAA are completely asymptomatic until they rupture, a trait which lends itself to the high associated mortality rate. AAA are usually discovered by chance, during clinical investigations for other conditions (e.g. using ultrasound or plain film x-ray (Desai et al. 2010). While most AAA will remain asymptomatic, as the diameter increases pain and/or pulsating may be felt in the chest, abdomen, lower back or scrotum (Fauci 2008). Symptomatic aneurysms, as well as those identified as >55mm diameter (even if asymptomatic) are generally indicative of the need for surgical intervention. If ruptured, mortality rates as high as 90% have been observed, with over 75% of patients dying before surgical intervention can take place. For those patients who do receive surgical intervention post rupture, the mortality rate is still 46% (Greenhalgh and Powell 2008, Brown and Powell 1999). The anatomy of the aorta The aorta is the largest blood vessel in the body, and it is responsible for transporting oxygenated blood from the left ventricle of the heart directly to the rest of the body. Blood is ejected through the aortic valve into the ascending thoracic aorta and the aortic arch where it supplies the head and arms via the brachiocephalic trunk, subclavian and carotid arteries. The abdominal aorta begins approximately at the level of the diaphragm and ends at the iliac bifurcation (Rizzo 2015). The abdominal aorta supplies blood to the 48

63 kidneys via the renal arteries, and to the abdominal cavity through the mesenteric and coeliac arteries. The iliac arteries subdivide into the inferior and superior femoral arties to supply the lower limbs. The normal aortic diameter varies with age, sex and body habitus. In a study by Allison et al. (2008) in which 504 subjects underwent full body CT scans, the mean diameter of the aorta at the level just inferior to the mesenteric artery to be /- 2.9 mm, decreasing to /- 2.2 mm at the level of the iliac bifurcation. In the body there are two main types of artery: - Muscular: Generally smaller and located at the peripheries, display viscoelastic properties. - Elastic: Located closer to the heart with larger diameters, contain less muscle cells which display elastic properties. Regardless of type, arteries are comprised of three distinct layers, as shown in Figure 13: - The Intima is the inner most layer. In healthy arteries the intima may be as thin as a single layer of endothelial cells. - The Media forms the middle layer of the artery. The media is made up of a three dimensional network of muscle cells, elastin and collagen fibrils (Holzapfel and Ogden 2014). Fenestrated elastic laminae separate the media into several concentric, fibre reinforced layers. These layers are most numerous in vessels proximal to the heart, becoming less common in the distal, peripheral arteries. 49

64 - The Externa (more commonly referred to as the Adventitia) is the outermost layer of the artery. The adventitia is made up mainly of fibrous tissue, loosely surrounded by connective tissue (fascia). Wavy collagen fibrils, arranged helically, serve to reinforce the artery wall. Aortic disease originates from damage to the intimal layer by endothelial disruption, this in turn leads to atherosclerosis which contributes to the weakening of the vessel wall. Hypertension can further compound and accelerate the expansion of an aneurysm. Treatment options FIGURE 13: CROSS SECTION OF THE AORTA, FROM HOLZAPFEL ET AL. (2000) The treatment of AAA in the early stages of the disease progression are mainly prophylactic and aimed at reducing the risk of catastrophic rupture. Figure 14 below shows the management plan proposed by Sakalihasan et al. (2005). 50

65 FIGURE 14: MANAGEMENT PLAN FOR TREATMENT OF AAA, FROM SAKALIHASAN ET AL. (2005) Surgical intervention is generally not performed if the diameter of the AAA is less than 5cm unless there are confounding factors such as distal embolisation, urethral compression or a contained haemorrhage. Surgical Aneurysm Repair Open surgical repair is the gold standard of care for AAA patients due to high rates of technical success, excellent long term durability, and infrequent requirement for postoperative re-intervention(sakalihasan et al. 2005). Open surgical repair is an exceptionally invasive procedure. Access to the abdomen is usually gained via a long midline incision (from the sternum to just above the pubic bone) or alternatively via a wide transverse incision. After gaining access to the abdominal cavity the intestines etc. must be manipulated to facilitate clear visualisation of the aneurysm. The neck of the aneurysm must be identified in order to control it if the neck is suprarenal or intrarenal then cross clamping above the renal arteries may be required briefly. The iliac arteries are controlled in a similar fashion. The diseased section of the aorta is then replaced with a graft, usually polyester or PTFE sealed with collagen (Li and Kleinstreuer 2006). 51

66 Contraindications Due to the length of surgery and the requisite level of invasiveness, open repair poses serious health risks to patients with co-morbidities. Those patients who are elderly, suffering from congestive heart failure, valvular heart disease, arrhythmias, COPD, renal insufficiency or high anaesthetic risk may be deemed inoperable. Endovascular Aneurysm Repair (EVAR) EVAR is a minimally invasive procedure to treat AAA. First introduced in the late 1980 s, it has become widely used as an alternative to open surgery in patients who are contraindicated. The benefits of EVAR over open surgery are reduced post-operative complications, reduced blood transfusion, no requirement for CPB, reduced ITU and reduced time to discharge (Richards et al. 2009). During EVAR, a covered stent graft is placed along the length of the aneurysm, sealing above the neck and below the aortic bifurcation in the iliac arteries, thus creating a conduit for blood flow while preventing further ingress of blood into the aneurysm sack. The resulting reduction in pressure in the aneurysm dramatically reduces the risk of rupture and stops the growth of the aneurysm. EVAR is routinely performed in the OR with a portable C-arm fluoroscope as, unlike TAVI, there is a higher chance of conversion to open surgery (bailout procedure). Also, unlike TAVI, there are currently no retrograde devices available or planned. EVAR is performed via a femoral cut down to gain arterial access. An introducer sheath ranging from 14Fr to 24Fr is inserted into the artery facilitating vascular access. A delivery system is passed through the introducer, along a guidewire, and into position using fluoroscopy. Once in position the stent graft is deployed, both self-expanding Nitinol and balloon expandable SS systems are available (Desai et al. 2010). A secondary stent graft is generally required on the contralateral leg to provide a conduit between the 52

67 main graft and the contralateral iliac artery. Where contralateral graft placement is not possible (in cases with hostile anatomy or in emergency rupture cases where time is critical) an aorta-uni iliac device is used with flow to the contralateral leg restored by subsequent surgical femoro-femoral grafting. Additional grafts and cuffs may be used to extend the overall length or provide additional fixation. Fixation is achieved primarily by the radial force exerted by the graft on the aorta, additional features, such as barbs or hooks can help prevent migration of the graft, particularly in the region of the neck Aortic stenosis As the public health environment improves, rheumatic valvular disease has been declining in the developed world. Valvular heart diseases however, have shown no trend of decreasing with the number of degenerative valve diseases on the rise. In the population over 75 years of age, 12.5% have degenerative valvular disease with 33.9% suffering from Aortic Stenosis (Yuan et al. 2013). Open Surgical Aortic Valve Replacement (SAVR) is the only treatment that has been shown to improve symptoms, functional status, and survival (Bonow et al. 2006). SAVR has been the gold standard treatment for symptomatic AS for decades with approximately 67,500 SAVR procedures carried out annually in the United States of America (Clark et al. 2012). Many elderly patients however, are contraindicated for open surgery due to their existing comorbidities (Zajarias and Cribier 2009). The aortic valve The aortic valve is the most important of the four valves in the heart. It functions as a one way valve that opens and closes due to differences in pressure gradients on either side of the valve. When the heart is in systolic contraction, the increased pressure in the heart causes the aortic valve to open, allowing oxygen rich blood to be pumped into the aorta and out into the body. 53

68 During diastole the pressure in the aorta exceeds the pressure in the heart and the aortic valve closes, preventing blood rushing back into the heart. As the left ventricle is tasked with supplying oxygenated blood to the entire vascular system, it bears the heaviest workload in the body. The result of this workload is that the aortic valve is the most susceptible to acquired heart valve disease (Freeman and Otto 2005). Figure 15 shows the aortic valve is comprised normally of three leaflets (tricuspid), with two leaflets (bicuspid) less prevalent forming the most frequent congenital cardiovascular mal-formation in humans (Tadros et al. 2009). FIGURE 15: SCHEMATIC VIEW OF AORTA AND AORTIC VALVE In extremely rare cases the aortic valve may comprise only one leaflet (unicuspid). The normal aortic valve is formed from 3 layers, the ventricularis on the ventricular side, the fibrosa on the aortic side and a layer of loose connective tissue, the spongiosa, between the two (Freeman and Otto 2005). These layers work in tandem to provide the required tensile strength and pliability for decades of repetitive motion, with the average human heart going through approximately 2.5 trillion cardiac cycles over the course of 65 years. 54

69 Over the course of our lives however, the aortic valve is not immune to disease; the initiation of valvular disease, from sclerosis to stenosis is discussed below. Calcific aortic valve disease Calcific aortic valve disease is a disorder with a spectrum of severity ranging from mild valve thickening without obstruction of blood flow, termed aortic sclerosis, to severe calcification with impaired leaflet motion, or aortic stenosis. In the past, this process was thought to be degenerative due to wear-and-tear of the leaflets with passive calcium deposition. Clinical data now suggests that calcific valve disease is a progressive disease akin to atherosclerosis (disease of the arteries characterized by the deposition of fatty material on their inner walls, i.e. coronary artery disease), with lipid deposition, inflammation, and active leaflet calcification. The overlap in the clinical factors associated with calcific valve disease and atherosclerosis and the correlation between the severity of coronary artery and aortic valve calcification provide further support for a progressive disease (Freeman and Otto 2005). Aortic sclerosis physiology Studies of early aortic sclerosis show plaque-like lesions on the aortic side of the leaflet extending into the adjacent fibrosa layers. These early lesions can be initiated by endothelial disruption as a result of increased mechanical or decreased sheer stress similar to early atherosclerotic lesions. Mechanical stress is highest on the aortic side adjacent to the attachment of the leaflets to the aortic root. Sheers stress is lower on the noncoronary leaflet due to the lack of diastolic coronary flow which may explain why this leaflet often presents with lesions first (Freeman and Otto 2005). Patients with congenital bicuspid valves present up to 2 decades earlier than those with tricuspid valves (Beppu et al. 1993, Pachulski and Chan 1993). Aortic Sclerosis can be diagnosed using 55

70 echocardiography and typically presents as focal thickening at the centre of the leaflets with commissural sparing and normal leaflet mobility. Aortic stenosis physiology AS is the most common valvular disorder in older adults with a prevalence of approximately 8% in patients over the age of 85 (Ambler et al. 2005). It accounts for approximately 60% to 70% of all valve surgery in the aging population (Coeytaux et al. 2010). In adults with severe symptomatic AS the 2- year mortality is approximately 50% (Ross and Braunwald 1968) with a five year survival rate of <20% (Horstkotte and Loogen 1988). Typically the symptoms of AS include angina, syncope, and eventually lead to heart failure. Like aortic sclerosis, AS is diagnosed using echocardiography to evaluate thickening and calcification of the leaflets (Figure 16) including the commissural posts. Decreased valvular area, antegrade jet velocity (aortic regurgitation), and mean pressure gradient across the valve are also primary markers of AS. Studies have shown an average increase in antegrade velocity of 0.3 m/s, an increase in mean trans-aortic pressure of 7mmHg and a decrease in valvular area of 0.1cm 2 per year (Otto et al. 1989, Roger et al. 1990, Faggiano et al. 1996). The most common symptom of AS is decreased exercise tolerance owing to the restriction of blood flowing from the heart to the rest of the body due to the narrowing of the aortic valve. Ventricular hypertrophy is also common is patients with AS, as the heart tries to compensate for decreased stroke volume. FIGURE 16: HEAVILY CALCIFIED AORTIC VALVE 56

71 Treatment options for aortic stenosis Medication, Balloon Angioplasty (BA), and surgical intervention are considered the traditional progression of treatment for those suffering from AS, however medication and BA treat mainly symptoms rather than slowing or reversing the progression of the disease. Surgical Aortic Valve Repair (SAVR) SAVR remains the gold standard treatment for those suffering from severe symptomatic aortic stenosis (Schwarz et al. 1982, Yuan et al. 2013). Stenotic heart valves cannot be repaired once extensive calcification has occurred or if there is congenital malformation. Thus, surgical valve replacement with prosthesis is required. Initially mechanical heart valves were implanted beginning in the 1960 s, constructed from polymers, stainless steel (or other alloys such as cobalt chromium/titanium) and a sewing ring formed from various fabrics e.g. nylon. From the 1980 s onwards, mechanical bi-leaflet valves were the most frequently implanted prosthesis to treat aortic valve disease. Biological tissue valves today are created from tissue (usually pericardium) harvested from bovine, porcine or equine sources mounted on a stainless steel or Nitinol frame, again with a sewing ring for attachment. Open SAVR is usually performed via a full median sternotomy and using CPB (Cohn 2008). With a patient on CPB the aorta is cross clamped, transacted and the native valve excised. An appropriate mechanical or stent-mounted tissue valve is then sewn in place of the diseased valve, the aorta is repaired, de-aired and the patient taken off CPB. Open SAVR is a highly invasive procedure resulting in a lengthy recovery time, however, with correct patient selection SAVR has a low operative mortality rate (Schwarz et al. 1982). Contraindications for SAVR Approximately 30% of patients suffering from severe AS are contraindicated for SAVR due to multiple comorbidities, frailty, porcelain aorta, poor 57

72 pulmonary function or left ventricular dysfunction (Iung et al. 2005). This cohort of patients have driven the need for less invasive surgical techniques as well as the inherent benefits such as reduced postoperative pain, faster postoperative recovery and shorter ICU stay (Cohn 2008). Transcatheter Aortic Valve Implantation (TAVI) Transcatheter Aortic Valve Implantation (TAVI) is a rapidly growing field of minimally invasive procedures as an alternative to standard surgical treatment in elderly or high risk patients with severe AS. Whereas standard SAVR requires a full sternotomy with a recovery time of up to 18 months, TAVI requires only a small cut down of the femoral artery or mini anterolateral thoracotomy. Furthermore TAVI can be performed as a beating heart procedure with no requirement to place the patient on Cardio Pulmonary Bypass (CPB) (Walther et al. 2009). TAVI is performed predominantly under fluoroscopy in traditional or hybrid operating theatres. As with any new technology there is a steep learning curve and relatively high adverse event rates, particularly in new centres. Figure 17 shows the difference in mortality rates between the first and second half of patients undergoing TAVI in a new centre (13.3% and 5.9% respectively, n=270) (Gurvitch et al. 2011). FIGURE 17: 30 DAY MORTALITY RATES IN TAVI. FIRST 135 PATIENTS AND SECOND 135 (GURVITCH ET AL. 2011) 58

73 The principal advantages of this percutaneous approach are: - No sternotomy required, greatly reducing recovery times. - No Cardiopulmonary Bypass required. - Beating heart surgery. - Greatly reduced procedural time. - The advantages of transapical over transfemoral are that, while more invasive, it offers an alternative for patients with peripheral artery disease, tortuous vasculature, or previous vascular complications (Singh et al. 2008). In any endovascular procedure, treatment can be delivered in two directions; antegrade with the direction of the blood flow, and retrograde against the direction of blood flow. In procedures involving the arterial system any device advanced from a peripheral location towards the heart is termed retrograde due to the direction of blood flow from the right heart to the peripheries and vice versa for the venous system. Retrograde systems Transfemoral TAVI systems (Corevalve, Evolute, Lotus, Accurate TF, Portico) gain procedural access via the femoral artery. An arterial puncture is created with a typical Seldinger technique and an appropriate sized introduce sheath is inserted (usually in the order of French size). A super stiff guide wire is introduced and tracked into the left ventricle (such as the Amplatz Super Stiff or Safari pre-shaped, both Boston Scientific). The TAVI delivery system is then tracked through the iliac arteries, the abdominal aorta, over the aortic arch, and crossing the Aortic valve to enter the left ventricle. 59

74 Antegrade Systems Transapical TAVI systems (Sapien, Engager, Accurate TA) use a left anterolateral mini-thoracotomy and pericardiotomy to visualise the apex of the heart. The surgical apex of the left ventricle is identified and punctured with a Seldinger type needle to introduce a guide wire across the native valve (Walther et al. 2009). The replacement valve is positioned within the native valve under fluoroscopy by advancing the delivery system over the guide wire. This is performed under rapid ventricular pacing (RVP) of beats/min to minimise transvalvular flown reducing the risk of valve migration during deployment (Rodés-Cabau et al. 2010). Once correctly positioned, the Aortic Valve Prosthesis (AVP) is deployed to its full diameter, affixing it against the aortic wall and ventricle. A summary of TAVI systems and implantation methods are shown in Table 6. 60

75 TABLE 6: TAVI DELIVERY SYSTEMS Device name Company Delivery Operating modality Corevalve & Evolute Medtronic TF, TSc, Dao Single control Knob, quick release slider Engager Medtronic TA Rotational Knobs, 2 stage deployment with safety button before stage 2 Acurate TA Symetis TA, Rotational Knobs, 2 stage deployment with safety button before stage 2 Acurate TF Symetis TF, Tao Rotational Knobs, 2 stage deployment with safety button before stage 2 Jennaclip JennaValve TA Cathlete plus: Single Control Knob, 3 stage deployment with safety button activation before each stage. Portico SJM TF Deployment wheel, dual locking buttons, 80% safety release button, slider. Lotus Boston TF Single rotational Knob, slider acts as saftey button before release TF = Transfemoral. TA = Transapical. TAo = Trans Aortic. DAo = Direct Aortic. TSc = Trans Subclavian. 61

76 2.6. Difficulties with endovascular delivery system use To understand the complexities involved in an endovascular procedure a task analysis of TAVI systems was undertaken, as discussed in section two different approaches are commercially available in delivery systems; antegrade and retrograde. Retrograde: Percutaneous access is achieved and an appropriate sized introducer placed in the femoral artery, a super stiff guidewire is advanced over the aortic arch, across the native valve and placed in the left ventricle, a pooled task analysis for several devices is shown in Table 7. TABLE 7: RETROGRADE TAVI DELIVERY SYSTEM TASK ANALYSIS Pooled Transfemoral Delivery System Deployment step DS is tracked over the guidewire through the introducer, over the aortic arch and across the native valve. Acurate TF Corevalve & Evolute * * * Portico Initial position is checked. * * * * Lotus Deployment Stage 1: Control knob 1 is CCW CCW CCW CCW rotated to begin deployment of inflow portion of valve in LVOT. Position is evaluated under fluoroscopy, * * * * valve can be re captured prior to safety button activation 2 Safety button depressed, removed. * * * Deployment stage 2: Control knob 3 is rotated CCW CCW CCW to release outflow portion of the valve in the aortic root. Pull DS through valve (to aortic side) close, * * * * and withdraw. Close access site to achieve haemostasis. * * * * CCW = Counter Clock Wise, CW = Clock Wise *= Step included for this device 1 Acurate TF uses distal control knob. 2 CoreValve has no safety button, can only recapture up to ~1/3 deployed 3 Acurate TF uses proximal control knob 62

77 Antegrade: A mini anterolateral thoracotomy is performed between the 5 th and 6 th intercostal space, the pericardium is opened and the surgical apex of the left ventricle is established. A series of pledgleted purse string sutures are placed on the apex and it is pierced with an 18 gauge needle. A super stiff guidewire is advanced across the native valve and placed in the descending aorta, a pooled task analysis for several devices is shown in Table 8. TABLE 8: ANTEGRADE TAVI DELIVERY SYSTEM TASK ANALYSIS Pooled Transapical Delivery System Deployment step Acurate TA Engager Jennaclip DS is tracked over wire, dilating the myocardium * * of the left ventricle and across the native valve. Initial position is checked. * * * Deployment stage 1: Control Knob is rotated to CCW CW CW deploy outflow portion of valve in aortic root. Prosthesis is seated into the native * * * cusps/annulus. The valve can be recaptured before safety button * * * activation. Safety button pressed/removed * * * Deployment stage 2: Control knob is rotated to CCW release inflow portion of valve Deployment stage 2.5: Control knob is rotated to CW CW release outflow portion of valve Deployment stage 3: second control knob rotated CW to release inflow portion of the valve Push DS through valve (to aortic side) close, and * * * withdraw. Close purse string sutures to achieve haemostasis. * * * CCW = Counter Clock Wise, CW = Clock Wise *= Step included for this device 63

78 As discussed in section 2.1.3, implantation depth of an aortic valve prosthesis in the LVOT is the only significant predictor of PPI. The difficulty in implanting a TAVI valve lies in the length of the delivery system. For transfemoral delivery systems, working lengths of 1200mm are common (as shown in Figure 18 below). Taking the reported difference in depth of implantation that resulted in new conduction disturbances/ppi as ~3.6mm (Binder et al. 2013), that represents a margin of error of just 0.3% of the total working length of the device. As such, even the most minimal influence of the delivery system design on implantation has the potential to result in an adverse event. Similarly, the landing zone during implantation of a stent graft to treat AAA can be as little as 10mm between the renal arteries and the neck of the aneurysm, with delivery systems of a similar length (up to 1500mm for thoracic aneurysm devices). In order to design the most robust delivery systems for endovascular interventions, as well as adhering to the applicable standards, it is necessary to perform usability tests throughout the design process as discussed previously. Mitigating delivery system related error, in as much as is possible, is the ultimate goal of usability testing. FIGURE 18: MEDTRONIC COREVALVE EVOLUTE DELIVERY SYSTEM 64

79 2.7. Summary of literature review Adverse events in endovascular treatment of AS and AAA are significant, particularly in new centres. Human factors (usability) testing is an integral part of the regulatory requirements to develop robust and safe medical devices. Simulation as an integral part of usability testing of medical devices is prescribed by American and European standards. Simulation requires both functional and structural fidelity to facilitate a suspension of disbelief during usability testing. Hollow silicone anatomical models lack functional fidelity when evaluating endovascular delivery systems Research objectives The objective of the current research is: To develop a simulator that is suitable for formative and summative usability testing of endovascular medical devices that provides high functional fidelity and a sufficient level of structural fidelity to facilitate a suspension of disbelief during testing. To achieve this, several research objectives were identified under the headings of functional (1-3) and structural (4 and 5) fidelity: 1. A need for digital compound anatomical models which include calcifications and thrombus for usability testing of endovascular devices. The lack of ILT and calcifications has been shown to adversely affect aneurysm behaviour in computational models, however they remain absent from silicone anatomical models used in usability testing of medical devices. 2. A need to integrate rigid connectors into anatomical models for use in pulsatile vascular simulators. 65

80 Silicone anatomical models are normally installed in pulsatile simulators using cable ties, jubilee clips and/or PTFE tape. There is a need to develop anatomical models that can be readily interchanged in simulators for device testing. 3. A need to produce compound anatomical models with multiple material properties suitable for use in pulsatile vascular simulators. Traditional methods of manufacturing silicone anatomical models are not suited to producing multiple material models. This limits the functional fidelity of silicone anatomical models, negatively impacting usability testing. 4. A need to develop a portable usability test bed for use with compound anatomical models. In order to use compound anatomical models for usability testing, a test bed is required. The test bed should provide pulsatile flow, physiological pressure, and temperature for evaluating devices in use conditions. Expert users of endovascular devices are healthcare professionals, often consultants, who have a multitude of commitments that place time demands on them. A usability test bed that is portable is desirable to allow device testing on location. 5. A need for haemodynamic monitoring and fluoroscopic imaging simulation as part of the usability test bed. Suspension of disbelief is an important factor during usability testing of medical devices. Operators should experience sufficient levels structural fidelity during testing to facilitate this, for this simulated fluoroscopic imaging and real time haemodynamic monitoring is required. The following chapters will address the research objectives defined above. Chapter three details objectives one and two, Chapter four details objective three and Chapter 5 details objectives four and five. A proof of concept usability test was carried out and is described at the end of chapter 5. 66

81 Chapter 3: Segmentation of compound anatomical models 3.1. Introduction and research objective The first research objective identified a need for compound anatomical models. Compound anatomical models differ from hollow silicone anatomical models in that they include thrombus and calcifications. These additional components will provide a more representative model of the anatomy for use during device testing. While efforts have been made to segment compound anatomical models in the literature for CFD/FEA computer simulations (Wang et al. 2002), the novelty of this work is the segmentation of compound anatomical models specifically for physical recreation using multiple material 3D printing. To create a compound model, it is necessary to segment the individual structures from medical imaging. Anatomical segmentation is the process of isolating anatomical structures from medical imaging in order to produce three dimensional digital models. These models can then be exported in.stl file format and physically recreated by various means such as machining, 3D printing, and lost wax casting for the production of silicone anatomical models. The limitations of current silicone anatomical models are discussed in section For the purposes of this research, anonymised patient data was used in order to create real world anatomical models, thus providing the most realistic conditions during usability testing. The research question explored in this chapter is whether compound anatomical models can be segmented that are suitable for multi-material 3D printing, for use in pulsatile simulators (i.e. models that contain calcifications, thrombus and the aortic wall). The approach taken was to segment anatomical structures from patient CT scans using the Mimics and 3Matic software programs, then to explore and develop additional protocols for use with multi-material 3D printing. 67

82 3.2. Process for segmenting anatomical models The process of converting medical imaging into functional, 3D printed parts has three high level steps, as shown in Figure 19 (Rengier et al. 2010). 1) DICOM images are captured from various imaging modalities. 2) The images are segmented and post processed to prepare files suitable for printing 3) The anatomical models are printed using a 3D printer. STEP 1 STEP 2 STEP 3 FIGURE 19: THE PROCESS CHAIN OF MEDICAL IMAGE TO 3D PRINTED MODEL (RENGIER ET AL. 2010) For this work, anonymised CT scans of three AAA patients and one TAVI patient were used. Each scan was first imported into the Mimics software package (Research Version 17, Materialise, Belgium). Mimics is a software interface and image segmentation system for the transfer of DICOM information from a variety of modalities into an output file. It is also used for simulating, measuring and modelling biomedical applications (e.g. simulating stress loads on prosthetic knee replacements). Mimics uses 2D imaging data to create 3D models of true to scale dimension (1:1). The software uses marching squares algorithms to threshold and segment regions of interest according to greyscale 68

83 values in the DICOM files based on Hounsfield Units (HU). Accurate segmentation requires a strong knowledge of anatomy and the ability to map 2D planar images to 3D. Once segmented, polylines are generated for the segmented anatomy and are exported into the sister program, 3Matic (Research Version 15, Materialise, Belgium) for further processing. In 3Matic, further operations such as Boolean additions/subtractions and smoothing can be applied at the users discretion to create multiple part assemblies. The mimics interface is shown in Figure 20 overleaf. Panel A) shows the coronal view, panel B) the sagittal view, and panel C) the transverse view. Panel D) is the 3D reconstruction of the anatomy in respect to the 3 planes, in this use of an AAA model. The basic method to segment individual anatomical structures is well described in the literature (O'Sullivan et al. 2015). A description of the workflow process for anatomical segmentation in Mimics is as follows: A Region Of Interest (ROI) is identified; a bounding box is placed around the extents of the required anatomical structure in three planes (coronal, sagittal and transverse). A pixel within the anatomical structure is selected which returns a HU value, the user then selects the upper and lower HU values to include in the 2D mask. The 2D mask of each data slice of the scan can be edited individually to include/exclude additional pixels as required. The 2D mask is converted to a 3D model automatically, this 3Dmodel can be edited within Mimics to remove floating pixels or noise. The models are exported as.stl files for post processing in the 3Matic software. This process was developed for this research to create compound anatomical models, and is discussed in the following section. 69

84 FIGURE 20: USER INTERFACE OF THE MIMICS PROGRAM 70

85 3.3. Segmentation of anatomical structure to create compound models The segmentation of both the AAA and TAVI models in this research were undertaken with a view to 3D printing the resulting compound models on the Objet Connex 500 System (section 4.3). 3D printing is advantageous over the traditional lost material casting method (section 2.4.4) as several materials can be printed together without the need for additional operations, thus drastically reducing the amount of time required to produce patient specific models. The size and complexity of models is not a factor when 3D printed (within the limits of the machine print bed) and the speed of production is expedited by weeks in some cases Segmentation approach for AAA and TAVI The patient AAA models were separated into four discreet parts: The true lumen, aneurysm sac, calcifications, and aortic wall. Figure 21 shows the pre-segmented sagittal view of Patient B. A large aneurysm is present (blue arrow), the neck starts at the level of L1/L2 continuing to the bifurcation at L4/L5. Calcification is clearly present at the margins of the sac. ILT is observed surrounding the true lumen with more presenting anteriorly. Specs of calcifications can be seen on the borders of the aneurysm sac. The following paragraphs detail the approach to segment the individual components of the compound model. FIGURE 21: SAGITTAL VIEW OF AAA PATIENT 71

86 - True Lumen The true lumen is the path of blood flow that remains inside the aneurysm sack. In Figure 22, the true lumen is shown in yellow. To segment the true lumen, a threshold is applied to the ROI to select contrast media which is highly radiolucent under CT. Using the pre-set parameters for bone works well as contrast has a similar radiolucency to bone. The true lumen in the final model will be negative space (empty). When printed, it will be filled with support material which will be removed afterwards. FIGURE 22: SAGITTAL VIEW WITH TRUE LUMEN MASK Aneurysm Sac Intraluminal thrombus is a pseudo tissue comprised of coagulated blood, and is found in most clinically significant AAA (Riveros et al. 2015) shown in purple in Figure 23. As the aneurysm sack extends, slow flowing currents build on the inner surfaces, eventually leading to clot formation. Segmentation of the aneurysm sack is a more manual operation, as the relative HU values are similar to surrounding tissues. Other artefacts such as collateral arteries must also be removed manually to isolate the AAA, as they share similar HU values as the thrombus. FIGURE 23: SAGITTAL VIEW WITH ANEURYSM SAC MASKED 72

87 - Calcifications Calcifications are hard plaques that formed on the aortic wall before or during aneurysm expansion (purple specs on the aneurysm sac margin, Figure 24). Inclusion of calcifications in AAA models is of importance as the rigid structure fundamentally alters the mechanical response of the aortic wall to stresses. Calcification may be dispersed radially around the outer boundary of the aneurysm, or encapsulated within the thrombus; this is caused by further delamination of the layers of the aorta and progressive expansion behind the plane that the plaque resides on. The calcification mask is exported as a separate.stl file. FIGURE 24: SAGITTAL VIEW WITH CALCIFICATIONS MASKED - Aortic wall The true lumen, thrombus and calcifications were combined to form a homogenous model of the negative space within the boundary of the aortic wall Figure 25. This homogenous model was the used to create a shell of 2mm thickness to segment the aortic wall. The shell is an offset of the outer boundary of the homogenous model. This thickness was chosen as it is widely published as appropriate in the literature (Doyle et al. 2008). 73 FIGURE 25: SAGITTAL VIEW WITH ALL PARTS COMBINED

88 Results of segmentation approach Figure 26 shows the completed AAA model in situ on the 3 planes of the patient CT scan. Figure 27 shows the same model imported into 3Matic. The true lumen, outer wall, ILT, and calcifications are clearly visible. A cross section is shown to demonstrate how the calcifications are embedded in the ILT. The individual.stl files for the three elements were then exported to Object Studio software for 3D printing FIGURE 26: 3D RECONSTRUCTION OF AN AAA SHOWN IN THE CONTEXT OF THE 2D CT SCAN. (section 4.3.3). FIGURE 27: COMPLETED SEGMENTATION OF AAA, LUMEN, CALCIFICATIONS AND ILT WITH CROSS SECTIONAL VIEW AND LOCATION INDICATOR 74

89 Segmented model modification for use in pulsatile simulators In addition to creating compound anatomical models, additional modifications can be carried out in the 3Matic software program. The second research objective highlighted the need for easier integration of anatomical models into pulsatile simulators. Modifications to anatomical models for use in pulsatile simulators are explored below: Step 1: Segmenting models In this section, an aortic model was segmented for usability testing of a TAVI delivery system. The model contained the left ventricle, aortic root, the major branches off the aorta, and terminated at a level inferior to the bifurcation of the superior femoral artery. The segmentation of the aortic true lumen in Mimics is the same as previously described in the AAA example. Figure 30 shows the masked aortic lumen in the sagittal plane. Figure 30 shows the same view with the 3D rendered true lumen. Step 2: Modification of segmented models At this point, the true lumen model was exported into the 3Matic program. The aortic lumen is wrapped and the wall offset to create a hollow model of the lumen with a 2mm wall thickness as previously described. The ends of the aortic branches were then cut perpendicular to the centre line of the lumen. Rigid ports were created on the cut surface to allow the soft vessels to connect to silicone tubing. These modifications allow 3D printed anatomies to integrate into pulsatile simulators and attached directly to silicone tubing without the need for adapters, clips or sealing tape. This makes the installation, use and changing of different anatomical models more efficient when used as a training or experimental aid. Figure 30 shows the hollow aortic model with rigid ports on the same sagittal slice of the patient CT scan. 75

90 FIGURE 30: SAGITTAL VIEW WITH TRUE LUMEN MASKED FIGURE 30: SAGITTAL VIEW WITH 3D RECONSTRUCTION OF TRUE LUMEN FIGURE 30: SAGITTAL VIEW WITH POST PROCESSED AORTIC MODEL The workflow for creating modified anatomical models with rigid ports is described here:.stl files imported into work space from Mimics. Smoothing is applied to models according to guidelines provided in the Mimics Innovation Training. Anatomy is wrapped a process where models are made watertight by closing any sharp edges left by the triangular meshing of the model. The wrapped surface is offset to create a hollow part with a wall thickness prescribed by the user. The ends of the vascular model are cut perpendicular to the central axis. Additional elements such as bases, stands, or ports are added like a traditional CAD programme. Finished models are exported as multiple.stl files ready for 3D printing. 76

91 Figure 31 Shows a more detailed view of the aortic model before and after post processing in 3Matic. The additional rigid ports are shown in opaqque colours for clarity. In the bottom figure, the large bore purple port in the left ventricle is used as an input for a pumping mechanism connected to this port when used in a simulator. The different sections (each individual rigid port 11 in total, and the hollow aorta) were exported as separate.stl files for 3D printing, as discussed in the next chapter. FIGURE 31: TOP: INTERNAL LUMEN (CONTRAST PATH) OF THE LEFT VENTRICLE AND AORTA. BOTTOM: HOLLOW, 3D PRINTER READY FILE WITH RIGID CONNECTION PORTS 77

92 3.4. Summary of outcomes from this chapter In this chapter it was demonstrated that compound anatomical models can be segmented from CT scans of AAA patients. The process is more labour intensive than creating a hollow aortic model, however the resulting models provide a more realistic testing environment when combined with 3D printing as outlined in the next chapter. This in turn aids usability testing by providing higher functional fidelity to the operator. The modification of compound anatomical models to integrate connection ports is a significant step forward compared to traditional methods of manufacturing hollow silicone anatomical models. Compound anatomical models can be segmented with a view to 3D printing on multi-material printers for usability testing of medical devices. Modification of compound anatomical models to integrate rigid connection ports are an advantage of this approach over that of traditional manufacturing methods. 78

93 Chapter 4: 3D printing of anatomical models for usability testing 4.1. Introduction and research objective The third research objective identified a need to take the digital compound anatomical models created in chapter 3 and manufacture them to include multiple material properties producing compound anatomical models. The limitations of manufacturing compound models using silicone has been discussed in section Three Dimensional (3D) printing, also known as additive manufacturing or rapid prototyping, was first conceived of in the early 1980 s. Oxford defines it as; the action or process of making a physical object from a three-dimensional digital model, typically by laying down many thin layers of a material in succession (Stevenson 2010). Charles Hull was granted the first patent for stereolithography in 1986 and continued to author numerous patents in the area. He also developed the.stl file format, which would complete the electronic handshake between computer aided design software, and transmit files for the printing of 3D objects (Gross et al. 2014). The objective of this chapter is to address research objective three: To 3D print a multi-material compound anatomical for use in usability testing of medical devices D Printing technologies for anatomical models 3D printing has developed from its origins as an academic pursuit to the point where desktop consumer models are available for under D printing in the medical sector has gained traction at a rapid pace, from dental implants, custom orthopaedic cutting guides, and pre-operative planning through to microfluidic applications and biological scaffolds. A brief expansion of the processes and limitations of each 3D printing technology is presented here: 79

94 Stereolithography (SLA) SLA printing works by photo curing acrylic or epoxy based resins using a high intensity UV light source. A moveable platform sits inside a vat of liquid photoreactive resin where the UV light source traces a 2D cross section onto it, polymerising the resin upon illumination. Once completed, the platform descends a set amount (usually in the range of 25μm) and the UV beam traces another layer on top of the first. This continues until the model is fully completed. While SLA remains the most accurate of all 3D printing technologies, it is limited by post processing requirements, high material cost and waste, and the fact that only one material can be used in a single build Selected Laser Sintering (SLS) SLS printing uses fine grade polymer or metallic powder which are thermally fused together with a laser. A roller spreads a thin layer of powder (~100μm) across a movable bed and the powder is heated to just below the melting point of the material. A high powered laser then traces a 2D cross section onto the powder, fusing it together. The bed then drops down and another layer of powder is deposited on top. The process repeats until the model is complete. SLS is a more versatile methodology than SLA, as a wide range of polymer, ceramic and metallic powders can be used. It also produces some of the most structurally robust parts. SLS is, however, limited in terms of surface finish due to the uncured powder acting as the support material. Models are also prone to shrinkage as the material must be kept at a high temperature throughout the manufacturing process Fused Deposition Modelling (FDM) FDM is possibly the most widely used 3D printing technologies available today. FDM works by extruding thermoplastics and depositing each semi-molten layer on top of each other to produce a 3D model. The layer height in FDM can 80

95 range from ~ μm and has a notable staircase effect on the outer surface of parts. In the medical field, FDM can be used with bio-absorbable or biodegradable materials (e.g. PLA) to manufacture absorbable stents and tissue scaffolds. FDM can use multiple materials in each build but requires a separate extruder head for each material, this usually limits a machine to 2 or 3 materials Inkjet printing Inkjet 3D printing is predominantly powder based, similar to SLS. Inkjet printing uses a liquid binder deposited onto layers of powdered material to create 3D parts. The quality of the builds are less than that of SLS or SLA as the layer height is relatively high (200μm), and the potential for the liquid binder to bleed is high Laminated Object Manufacturing (LOM) LOM creates 3D models by stacking individually cut sheets of paper, plastic or metal on top of each other. A laser or razor cuts each layer into a 2D slice of the model and then glues or fuses each layer together. LOM can create quiet detailed models but is a far slower process than the others described Polyjet printing technology Polyjet printing is currently the most advanced 3D printing platform available. Polyjet printing works somewhat like a traditional inkjet printer in that two or more materials can be combined to create a multitude of secondary materials. The polyjet system is also unique in that it can combine soft, rubber like materials and rigid materials to produce over molds, flexible parts and gaskets. The current state of art for 3D printing is the Objet Connex 3 (Stratasys Ltd.); it has three colour cartridges and can create 3D parts in full colour, as in Figure 32, without the need for post processing or painting. 81

96 A summary of current 3D printing technologies is shown below in Table 9 (Rengier et al. 2010) TABLE 9: OVERVIEW OF ESTABLISHED 3D PRINTING TECHNIQUES USED IN THE MEDICAL ARENA (RENGIER ET AL. 2010) Technology Accuracy Cost Advantages Disadvantages Stereolithography (SLA) +++ $$ Large part size Moderate strength Selective Laser Sintering (SLS) ++ $$$ Large part size, variety of materials, good strength High cost, powdery surface Fused Deposition Modeling ++ $ Low cost, good strength Low speed (FDM) Laminated Object + $ Low cost, large part size Limited materials Manufacturing (LOM) Moderate strength Inkjet printing techniques + $ Low cost, high speed, multimaterial capability Most 3D printing technologies are only capable of using a single material to create anatomical model. To create compound anatomical models for this research, polyjet printing was selected as it is the most suitable technology. The ability to print several materials in a single model at the same time is ideally suited to this application. A more detailed over view of the Connex system is provided in the next section. FIGURE 32: FULL COLOUR HELMETS PRINTED ON AN OBJET CONNEX 3, STRATASYS LTD. 82

97 4.3. Polyjet printing using Objet Connex 500 The Dept. of Design and Manufacturing Technology (DMT) and the Materials and Surface Science Institute (MSSI) procured an Objet Connex 500 3D printer. The Connex system was chosen specifically for the ability to print flexible materials, and the potential biomedical applications inherent with that ability. The Connex 500 is a polyjet 3D printer that uses two base materials (acrylic monomer based resin) and a water soluble support material that are cured using high intensity UV light. The two base materials can be combined to create 14 digital materials with varying mechanical properties, from soft rubber analogues (~27A shore hardness) through to rigid materials (~85D shore hardness). The resolution of the printer is 600 DPI and the layer height (Z axis) is 30μm. The stated tolerance for the Connex system is +/- 0.2mm, but in practice, tolerances of +/- 0.05mm are achievable with careful cleaning of the parts. The build envelope is approximately 500 x 400 x 200 mm, allowing aortic models to be printed in full scale. As shown in Figure 33 an aortic model was printed from the bifurcation of the brachiocephalic trunk to the below the bifurcation of the superior femoral arteries in a single part. The model is shown on the Connex 500 print bed before the support material was removed. FIGURE 33: COMPLETE AORTA PRINTED ON THE CONNEX

98 Materials Historically, aortic models for use in pulsatile flow loops were made using silicone and the lost wax method (Doyle et al. 2008). This is a labour intensive method that requires the machining of pairs of aluminium molds per anatomical model, and significant time to create a wax core, mount the core in a secondary mold and fill the cavity between the two with silicone. Even then, there is no way of ascertaining if a model has any defects such as air bubbles or uneven wall thickness until after the silicone is fully cured and the molds separated. 3D printing eliminates this uncertainty and is significantly more time efficient. The Connex systems creates up to 14 digital materials from the two base materials of TangoPlus (rubber like) and VeroWhite (Rigid). The digital materials range from polypropylene simulants through varying stiffness s of rubber. Work has already been undertaken to characterise these 14 digital materials by researchers in the MSSI on campus. Cloonan et al. (2014) determined that the TangoPlus (FLX_930) material was similar to recognised silicone aortic analogues. A limitation of this work however, was that the materials were tested at room temperature. The materials printed on the Connex 3D printer are sensitive to temperatures with rigid materials entering glass transition (Tg) at approximately C. The effect of increased temperature on the rubber like materials is notable, subjecting the materials to body temperature (37 C) causes a dramatic reduction in shore hardness in practice. For this reason a harder material may be more suited as an analogue. Testing is currently underway in the University by Ms. Bronadh Lynch (MSSI) to characterise Connex materials against healthy and diseased aortic tissues in real world conditions. Early work suggests DM_9760 may be more suitable as an aortic analogue (a digital material with a shore hardness of 60D). For the 84

99 purposes of this research DM_9760 will be used to 3D print the aortic wall, TangoPlus for the thrombus and VeroWhite as the calcifications Process of 3D printing compound anatomical models The process of 3D printing anatomical structures using the Connex 500 is described as follows: All parts of the assembly are exported from Mimics in.stl format The parts are inserted as an assembly onto a build tray in the Objet Studio Software package (Figure 34). Material properties are applied to individual parts. The build is automatically checked for faults, overlapping parts and overall size. The printer is started, automatically pre warms and purges all print heads before beginning to print the model. The finished model is removed from the printer, support material is removed by hand and residue is removed using pressurised water. FIGURE 34: AAA MODEL IN THE OBJECT STUDIO SOFTWARE 85

100 Initial experience with compound anatomical models Initial experiments highlighted an important issue with the printer capabilities. In the first experiment, the discreet parts of the AAA were segmented as described in the previous chapter, however after printing the model it was found to have been printed as a single solid part (Figure 35). This part consisted of DM_9760, the material specified for the aortic wall. After investigation it was discovered that the.stl files of the individual parts overlapped, which caused the printer to register the model as a solid, applying the material properties of the outer surface (aortic wall) to the entire model. To overcome this issue, all parts of the model were imported back into 3Matic and the following process was derived: The aortic wall was subtracted from the ILT Calcifications were subtracted from both the aortic wall and ILT. The true lumen was subtracted from the ILT, aortic wall and calcifications. This results in a model where each part exists in the negative space of itself. A second experimental print was undertaken on a 10mm cross section of the FIGURE 35: PRINTED AAA WITH SECTION REMOVED SHOWING PRINTING DEFECT 86

101 aneurysm. The cross section is shown in Figure 36; left is a digital view of the aneurysm cross section in the 3Matic program, middle is a highlighted view showing the location of the cross section in relation to the AAA, while right shows the 3D printed version of the section. The boundaries of the Aortic wall, ILT and Calcifications can be clearly distinguished in the printed model. FIGURE 36: CROSS SECTION OF AAA SHOWING AORTIC WALL, THROMBUS, CALCIFICATIONS, AND TRUE LUMEN. Another test was carried out on a 10mm cross section just superior to the femoral artery bifurcation. Figure 37 shows the arterial wall printed in a flexible material with the heavy calcifications printed in a rigid material. Material selection is discussed further in section FIGURE 37: CROSS SECTION OF ILIAC ARTERY SHOWING CALCIFICATIONS 87

102 Experiments were also undertaken to evaluate the modification of anatomical models. An important question was whether the interface between the rigid additions and the flexible anatomical models would be secure and water tight when used in pulsatile flow simulators. Figure 38 shows the result of the initial test print. The border between the rigid and flexible materials can be clearly observed. The interface was found to be well connected and no mechanical weaknesses observed. Figure 39 demonstrates how the rigid ports are used in practice. A Large bore silicone tube (30mm internal diameter) is easily and securely connected to the rigid port on the left ventricle with an interference fit. FIGURE 38: LEFT, A DIGITAL VIEW OF THE HOLLOW ANATOMICAL MODEL WITH RIGID PORTS ATTACHED, RIGHT, THE 3D PRINTED RESULT FIGURE 39: 3D PRINTED LEFT VENTRICLE, RIGID PORT AND SILICONE TUBING 88

103 Results of 3D printed compound models The models created with the Connex 500 are more dimensionally accurate, considerably faster to produce than silicone versions and incorporate multiple materials. Depending on the height of the model (in the Z axis) then models can be produced over night. Figure 40 shows two AAA models including thrombus and calcifications; the top model took approx. 12 hours to print while the bottom model took approx. 16 hours due to the larger [taller] aneurysm sack. Figure 41 (overleaf) shows the Mimics models of an AAA, the finished model on the bed of the Connex 500 printer, and the model after cleaning. FIGURE 40: ADDITIONAL AAA MODELS PRINTED INCLUDING CALCIFICATIONS 89

104 FIGURE 41: (A) COMPUTER MODEL OF AAA, (B) 3D PRINTED MODEL ON THE PRINTER BED, (C) 3D PRINTED MODEL AFTER CLEANING. 90

105 4.4. Summary of outcomes from this chapter The Connex 500 system is ideally suited to producing compound anatomical models with multiple material properties assigned to individual elements. The method for creating compound models as discussed in the previous chapter was not without issue. The initial printing experiments demonstrated issues in the preparation and segmentation of the files that caused printing failures. The modified protocol eliminated these issues with three AAA models and a full sized aortic model successfully printed. Compared to the current gold standard, the process of applying the additional protocol and printing the finished models was completed within 24 hours this represents a time saving of several days (or weeks) per model and is less than the curing time for most silicones. One disadvantage of 3D printing anatomical models is the removal of support material from inside the models. Wax cores used for silicone can simply be melted to remove supports. The Connex system support material must be removed by hand before residue is removed with a high pressure water jet. Overall the quality and functionality of 3D printed models is worth this minor drawback. Multiple material compound anatomical models can be printed on the Connex 500 system. Additional protocols are required when segmenting compound anatomical models to ensure that all elements print correctly. 3D printing compound anatomical models offer increased functional fidelity and a significant time saving when compared to hollow silicone models. 91

106 Chapter 5: Development of a usability test bed for 3D printed compound anatomical models 5.1. Introduction and research objective The fourth research question raised the need for a usability test bed compatible with compound anatomical models. Any test bed for use with endovascular devices should mimic physiological conditions as closely as possible. Pulsatile flow, physiological pressure and temperature are essential for representative testing of implants and delivery systems alike. The proposed usability test bed is designed to accept multiple anatomical models and disease states. The addition of standard sized rigid ports to anatomical models discussed in Chapter 4 makes the models readily interchangeable in the simulator. The objective also states the importance of portability. Lab based simulators require expert users to commute to perform device testing. This may be appropriate at the summative stage of testing when formal acceptance criteria are required, however this is impractical during formative testing, particularly if there are multiple iterations of the design. The fifth objective identified the requirement for real time haemodynamic monitoring and simulated fluoroscopy imaging during simulated use testing. These elements lend themselves to the suspension of disbelieve which is crucial for valid simulated use by operators. The objective of this chapter is to address research needs four, five and six; to create a portable, self-contained pulsatile simulator with integrated imaging and haemodynamic monitoring capabilities for use with compound anatomical models. 92

107 5.2. Mechanical development A system level diagram for the simulator is shown in Figure 42. The compound anatomical model is fed by a piston pump which draws from a heated reservoir. The flow loops from the pump, through the anatomical model, a pressure control valve, and back to the reservoir. The heart rate and cardiac output (rate of flow) is digitally controlled through a Virtual Instrument (VI) in the LabView program (National Instruments, USA). The LabView interface simultaneously displays real time pressure, flow, and temperature data as well as simulated ECG signal on a screen within the simulator. A second screen displays a simulated fluoroscopic image captured with a CCD camera and processed with a video converter unit. A Computer Aided Design (CAD) model of the proposed design, shown in Figure 43, was created in Solidworks (Dassault Systems, France). A description of the development of individual elements are described below. FIGURE 42: SCHEMATIC OF PULSATILE SIMULATOR 93

108 FIGURE 43: SOLIDWORKS MODEL OF PLANNED SIMULATOR Pulsatile flow Physiological pulsatile flow in the body is caused by the pumping action of the heart. Pulsatile flow is an essential component of any system to be used in the evaluation of endovascular devices in order to recreate the mechanical stresses found in the body. For this work, a reciprocating piston pump was used, based on the work of Morris et al. (2013). A steel, direct drive, rack and pinion (Module 1 teeth, Radionics) is powered using a 200 step (1.8 ) stepper motor (Radionics, ). The motor is controlled with a dedicated stepper motor controller (Geckodrive, G201X Step Motor Control). The controller is sent direction and pulse information from an Arduino Uno (Rev 3, Matissimo, Italy), the Arduino is sent high level information from the LabView VI via a LabView DAQ card (National Instruments, USB 6008). Two one-way valves are installed in opposite directions in the piston pump manifold to create unidirectional flow. The length of the piston stroke determines the volume of fluid displaced, creating the 94

109 required cardiac output. The cardiac output can be user adjusted by changing the controls on the LabView VI (section 5.2.6). The piston stroke length is calculated in real time by the Arduino code. The code divides the cardiac output (litres per minute,) by the heart rate (beats per minute), then computes the volume output required per stroke using the π.r 2.h formula where r is the radius of the piston pump barrel. Finally the number of steps required per stroke is calculated based on the drive gear diameter. As an example, the stroke length (h) for the parameters cardiac output = 6l/m and heart rate = 72bpm is computed as: (π * 22 2 ) / (6000 / 72 * 1000) = (h) This is then converted to the required number of steps using a division of 0.15 (equating to 200 steps per 30mm): 54.83(mm) / 0.15(steps per mm) = (steps per stroke) The number of steps are then passed from the Arduino to the motor controller to drive the pump and generate the pulsatile flow Pressure Correct systolic pressure is generated by the cardiac output (approx. 80ml/beat) while the diastolic pressure is a controlled using the viscoelastic properties of the system. The anatomical model, silicone tubing, and control valve (distal to the aortic bifurcations) generate the required back pressure. Each anatomical model requires adjustment of the control valve due to the difference in compliance and resistance in the system. The pressure in the simulator can be measured at multiple locations, the current design iteration allows up to four channels of pressure monitoring to be displayed in real time on the VI. Each channel is monitored using a male 95

110 leur fitting on an eight inch extension tube, connected to a pressure transducer. In this configuration, pressure may be monitored by connecting standard fluid management lines to the male leurs (e.g. the pressure at the introducer hub). The pressure transducers selected are the MPX5050GC7U (Freescale Semiconductor, Inc. USA), and have a range of mmhg. The transducers are ratio metric between 0V and 5V, the output to pressure differential is shown in Figure 44. The sensors are connected to analog-in lines on the DAQ card. FIGURE 44: OUTPUT TO PRESSURE DIFFERENTIAL, FREESCALE MPX5050GC7U Viscosity As the viscosity of blood differs substantially from water, a blood analogue was used in this work. A blood analog commonly used in the literature is a composition of glycerol and water (~44:56 ratio) with sodium chloride added to match the concentration in healthy human blood (~0.9%) (Yousif et al. 2010, Doyle et al. 2008, Morris et al. 2013) Temperature The temperature of the simulator is maintained at a constant 37 C (+/- 0.5 C). The water is heated in the reservoir using a 600 watt immersion bar heater. The heater is controlled with a 240V relay (5V switching) connected to a digital line 96

111 output pin on the DAQ card. The temperature is monitored with a thermistor placed at the lowest point of the tank. The thermistor is connected to an analog input pin on the DAQ card with a 10k Ohm resistor. The digital output pin is coded to high (relay on) while the temperature is <37 C and low (relay off) if the temperature is >37 C. This maintains the simulator at the correct physiological temperature Simulated imaging The simulated fluoroscopy imaging is presented by using a CCD camera, connected to a video converter (Model JM1936). The video converter is used to convert the S-video signal to VGA compatible signal for display on the second computer monitor in the simulator. The image is also converted into black and white and the contrast and brightness adjusted to best approximate the image quality of a fluoroscope. Figure 45 shows the on screen image presented to the operator during a simulated aortic root arteriogram before and after an injection of contrast (dye). FIGURE 45: SIMULATED FLUOROSCOPIC IMAGE ON THE SIMULATOR 97

112 LabView interface A Virtual Instrument (VI) was created in LabView as an interface to visualise both input parameters (pump controls) and output haematological measurements. The VI was designed to mimic a typical monitoring screen found in the cath lab. The VI allows adjustment of the input parameters and monitoring of the outputs in real time: Inputs - Beats per Minutes - Stroke volume - Rapid pace (~ 200bpm with negligable flow used when implanting TAVI valves) Outputs - Up to four simultaneous prssure measurements - Max/min pressure values - Temperature - Cardiac output (flow) - Heater status FIGURE 46: LABVIEW INTERFACE OF THE SIMULATOR 98

113 Electronic controls The control of the simulator is coded entirely from beginning on the LabView and Arduino platforms. LabView is a graphical programming environment where as Arduino uses a modified version of C++, the unabridged codes are contained in Appendix I and II respectively. The Arduino is sent information from LabView via the DAQ card to control the stepper motor. All other measurements and controls are routed through the DAQ card alone Housing and portability The simulator is built into a padded flight case for safety and portability. The case is constructed from 9mm birch plywood with a protective HDPE outer layer. The lid contains the two computer monitors to display imaging and VI information which lock at 90 when opened. The entire simulator can be transported by car and requires only a single electrical socket to run. 99

114 5.3. Results of portable usability test bed The completed simulator was built and tested using an idealised aortic model (Figure 47). Figure 48 shows an operator performing a balloon angioplasty of the right femoral artery. The completed simulator was presented at the Materialise World Conference 2015 in Brussels (O'Sullivan et al. 2015). A video demonstration of its functionality is presented on the accompanying CD. A pilot usability test was also performed out using the simulator, as discussed in the next section. FIGURE 47: FRONT VIEW OF THE COMPLETED SIMULATOR. FIGURE 48: LEFT, OPERATOR PERFORMING FEMORAL ARTERY ANGIOPLASTY, RIGHT, SIMULATED FLUOROSCOPY IMAGE ON DISPLAY DURING PROCEDURE 100

115 Introduction Proof of concept usability test using the completed simulator A Proof Of Concept (POC) usability test was conducted using 3D printed compound anatomical models in the completed simulator. The purpose of the POC test was to evaluate the peak force required to actuate a transcatheter aortic valve delivery system using different grip types and handle diameters. Two participants, one male and one female, both aged 22, took part in this POC test. Equipment A torque gauge was connected to a force meter (Mecmesin, AFG Mk4 1000n) in order to record Maximum Voluntary Exertion (MVE) of the participants. A T bar was attached to the torque meter to provide a rigid grip for testing MVE. EMG sensors were placed on the participants skin over the body of the Flexor Carpi Ulnaris (FCU) muscle according to the guidelines by Perotto et al. (2005) in order to collect Electromyography (EMG) data. The EMG signal was collected using the Nexus-10 MKII system (Mind Media BV, Herten, Netherlands). The collected data were exported as a.txt file from the accompanying BioTrace+ software, and processed in Microsoft Excel. Delivery system For the purposes of this POC test a generic delivery system was created in Solidworks, and printed using the Connex 500 system. The delivery system was designed to accept multiple handle and used the shafts and aortic valve from the CoreValve Evolute TAVI system (Medtronic Inc. Dublin, Ireland). Figure 49 shows the generic delivery system with a small diameter handle (20mm). FIGURE 49: 3D PRINTED GENERIC TAVI DELIVERY SYSTEM 101

116 Figure 50 shows the three different sized handle used in this test: 20mm, 40mm and 60mm respectively. Each of the handles was easily interchanged with the generic delivery system, providing a range of test conditions with a single delivery system. FIGURE 50: INTERCHANGEABLE HANDLE FOR GENERIC DELIVERY SYSTEM Procedure Each participant first had their skin prepared, before placing the EMG electrodes over the body of the FCU (Perotto et al. 2005). They were then asked to grasp the T bar in the torque meter and to rotate the bar anti clockwise with to their respective MVE. The EMG signal was recorded during this MVE run. The delivery system was loaded with a prosthetic aortic valve frame, and inserted into the completed simulator. For each run, the delivery system was placed in the simulator through a 10 French introducer sheath, tracked over a guidewire across the aortic arch into the aortic root, and left for two minutes before actuation. This allowed time for the delivery system shafts and nitinol stent frame to reach body temperature, and for the shafts to have to move through the aortic arch. This provided the most realistic test scenario to evaluate force during valve deployment. The participants then took it in turns to actuate each of the three handle sizes, with two different grips, pinch grip and power grip. During the task the participants were instructed to maintain the distal position of the delivery 102

117 system relative to a mark placed at the atrioventricular junction. Visualisation was only permitted on the integrated simulated imaging system. The participants An EMG recording was taken for each run. Pilot test results The POC usability testing of a TAVI delivery system was successfully completed using the newly developed simulator and 3D printed compound anatomical models. Due to the sample size it is impractical to draw any meaningful conclusions from the EMG data collected. The force data does suggest that as handle diameter increases, a higher peak force can be generated. This broadly supports the established literature (Weinger et al. 2010). The peak forces (Nm) for each grip type and handle diameters are shown in Figure 51. Each of the EMG recordings were converted to % MVE, a comparison of pinch grip between the two participants is shown in Figure 52 overleaf. Peak Force Generated (Nm) Nweton Meters (Nm) Handle Diameter (mm) Male Pinch Grip Male Power Grip Female Pinch Grip Female Power Grip FIGURE 51: PEAK FORCES GENERATED BY SEX AND GRIP TYPE 103

118 % MVE Male Pinch Small Nm % MVE Male Pinch Medium Nm % MVE Male Pinch Large Nm Axis Title Female Pinch Small Nm Axis Title Female Pinch Medium Nm Axis Title Female Pinch Large Nm FIGURE 52: INDIVIDUAL MAXIMUM VOLUNTARY EXERTION PLOTS FOR PILOT TEST 104

119 5.4. Summary of outcomes from this chapter In this chapter, a pulsatile flow simulator was developed for use with 3D printed compound anatomical models. The simulator features a real time haemodynamic monitoring system, including pressure, temperature and flow rate, displayed through a purpose built LabView VI on one of two built in screens. On the second screen, a simulated fluoroscopy image of the simulator was displayed. The entire simulator (pumping and heating mechanisms, anatomical models and dual screens) are contained in a flight case that is easily transportable for onsite usability testing with expert users. The design was built and tested using aortic and AAA models, with a TAVI valve successfully deployed in the full aortic model. A pilot usability test was carried out to investigate the maximum force generated while deploying a TAVI valve in the simulator. The pilot test used electromyography to record electrical activity in the flexor carpi ulnaris during valve deployment. The key advantage of the work presented in the flexibility to support multiple anatomical models and disease states. Flow rates can be easily adjusted to match that of aortic, peripheral or cerebrovascular models. A portable simulator for use with 3D printed compound anatomical models can be developed. Simulated real time haemodynamic monitoring can be achieved using a LabView VI. Simulated fluoroscopic imaging can be integrated into a portable simulator. 105

120 Chapter 6: Discussion Simulation during usability testing is a critical component of the design process for medical devices. Eliciting user needs in the early stage of the process and validating the resulting design against those needs creates a robust, safe and effective device. The primary aim of the current research was to develop a simulator that is suitable for formative and summative usability testing of endovascular medical devices. The simulator must provide high functional fidelity and a sufficient level of structural fidelity to facilitate a suspension of disbelief during testing. To achieve this aim, the research was comprised of five research objectives Objective one: Segmenting compound models The first research objective primarily addressed the need to digitally segment compound anatomical models i.e. models that include calcifications and thrombus. A protocol to digitally segment compound anatomical models was successfully developed and several patient models were created for multi material 3D printing. The protocol did however, require modification after initial test prints were carried out. This is discussed in objective three. Throughout the research, the Mimics Research software package was used. While other segmentation software solutions are available (i.e. Osirix, ITK- Snap), Mimics is specifically geared towards research applications. Other segmentation programs are primary aimed at medical professionals and preoperative planning (such as the measurement of AAA for grafts). Mimics research edition has the ability to modify the anatomical segmentation and export the resulting models to perform CFD/FEA in programs such as Abaqus. The quality of the DICOM scan images used with the Mimics program dictates the quality of the models produced. For example, the segmentation of the aortic wall is difficult for a number of reasons. The wall its self has similar 106

121 properties and density to surrounding tissues making it difficult to delineate. The scanner resolution and the relatively small size of the aortic wall causes further issues. The methods of creating the aortic wall in Mimics presented here is based on previous work in the literature (Cloonan et al. 2014, Doyle et al. 2008). A uniform offset is generated from the negative space of the true lumen. A general assumption used from the literature is a thickness of 2mm applied to silicone and computational models. In reality, the thickness off the aorta tapers distally from >2mm at the aortic root, to ~1.3mm at the level of the bifurcation in response to reducing pressure. A limitation of the Mimics package in this application is that offsets must be uniform throughout the model. In section 2.5.1, the formation of the aorta is discussed with the three layers that comprise the aorta (the intima, media and adventitia) each having distinct mechanical properties (Rizzo 2015). Segmentation and production of physical models traditionally use a single material throughout, based on mechanical testing of the aortic wall as a whole. Mechanical characteristics of the individual layers are difficult to quantify due to the practicalities of isolating each. Ideally, a three layered aortic wall could be segmented in Mimics using multiple offsets and exporting each layer as a separate.stl file. The impact on device usability testing of these limitations should be minimal, however, further research is warranted to quantify any negative effects. In experimental models, the lack of discreet layers may have a notable effect, particularly for migration testing and evaluation of luminal damage. The effect of isolating the aorta without additional supporting structures is also worth considering during usability testing. The mechanical characteristics of the aorta have been well described in isolation, but the aorta itself is constrained in the body by various organs, fascia, and accessory arteries. How a device reacts in these isolated anatomical models, which have a far greater range of movement than in clinical practice, needs to be explored. Several 107

122 important device evaluation criteria are missed when testing devices in hollow anatomical models, and this in turn can impact the validity of the simulated use testing. The reaction of smaller vessels (such as the femoral arteries) to straightening cannot be considered valid if the presence of calcifications are omitted in silicone models. The same is true of TAVI implantation. Paravalvular leak is a concern with every new valve design; however testing in a compliant silicone model will provide results skewed towards the positive. In AAA implantation, the angulation of the neck can vary significantly depending on the shape of the true lumen within the thrombus. This work aims to mitigate the possibility of false positives during device testing. While the work presented here on creating and modifying compound anatomical models is framed in the context of usability testing of medical devices, there are applications for the methods in both computational fluid dynamic research and experimental testing (such as migration and rupture studies) Objective two: Modifying anatomical models The second objective dealt with the modification of anatomical models. Section discussed current limitations of physically integrating silicone anatomical models into pulsatile simulators. Installing and interchanging silicone anatomical models is a time consuming task that requires cable ties, jubilee clips and/or PTFE tape as well as rigid couplers. As access to expert operators is often limited, the ability to efficiently interchange multiple models is desirable (such as 5 th and 95 th percentile tortuosity models). In cases where multiple iterations of a single device are required to be tested, the ability to replace an anatomical model that has a prosthesis deployed in it for an empty model offers a considerable opportunity to increase productivity. The work presented here demonstrates a method to include rigid ports to compound anatomical models. These ports are designed to press fit standard tubing without the need for additional components (Figure 53). 108

123 Further applications for modification of anatomical models include the ability to add restrictions to models. These restrictions could be applied to the rigid ports to generate a specific flow exiting an arterial branch. Alternatively structures such as plaques can be added to anatomical models derived from population data. These modifications can be customised by expert users to represent rare or challenging cases where scan data may not be readily available. An additional advantage of 3D printing rigid ports directly incorporated into anatomical models is that these models can be used in a Cath Lab or Hybrid OR during validation testing using a C-arm fluoroscope. As there are no metal connectors in the anatomical models and silicone hosing connecting the pump and model, there are no artefacts present under x-ray imaging. The ability to digitally modify anatomical models for physical manufacture of compound anatomical models is FIGURE 53: INSTALLED AORTIC MODEL WITH LEFT VENTRICULAR INPUT PORT. a key benefit of the work. 109

124 Objective three: 3D printing compound anatomical models The third research objective was to produce compound anatomical models with multiple material properties suitable for use in pulsatile vascular simulators. 3D printing in healthcare has become a useful tool for both visualising complex conditions and surgical planning (Rengier et al. 2010). The work presented here is a novel use of multi material in the healthcare setting. The production of compound anatomical models with multiple encapsulated materials is impossible with the current technology. Figure 54 demonstrates three clearly distinguishable materials contained within the one cross section. The calcifications (blue arrow) are dispersed within the thrombus (orange arrow) completely independent of the aortic wall (red arrow). This capability is unique to multi material printing. FIGURE 54: ENCAPSULATED MATERIALS Other advantages of 3D printing anatomical models over lost wax casting is the time and cost benefit. Due to the large size of the print bed on the Connex 500, three patient specific AAA models were manufactured in a single 26 hour build. While the polyjet technology is relatively expensive (each model costing approx ), this is similar to the cost of commercially available hollow silicone models. Furthermore, the time savings can be measured in days for each individual model. 110

125 There are some limitations to the 3D printed materials presented here. As with all 3D printed materials there are inherent weaknesses along the Z plane (parallel to the printer bed) caused by poor bonding between individual layers. The TangoPlus material is slightly hydroscopic and overtime can absorb some water in the surfaces in contact with water. This leads to opacification of the material, and in thin sections, can lead to failure as the material swells and delaminates (Figure 55). For this reason the use of pure TangoPlus for the aortic wall, as proposed by Cloonan et al. (2014), is inappropriate for prolonged use in contact with water. There are also issues with longevity of polyjet materials; the materials are UV cured and overtime can suffer from degradation due to exposure to light. Experiments are underway to investigate coating the finished 3D printed models in order to reduce this FIGURE 55: DELAMINATION ALONG THE Z PLANE (AORTA) phenomenon. The use of 3D printing is also Polyjet multi material printing represents a new paradigm for the production of anatomical models for usability testing of medical devices. While the work presented here is applicable to usability testing, use of this new approach could extend further into training, demonstration and experimental applications. 111

126 Objective four: developing a portable usability test bed Objective four described the need to develop a portable pulsatile flow simulator that accommodates 3D printed compound anatomical models. The design and use of pulsatile flow simulators is well discussed in the literature (Chong et al. 1998, Doyle et al. 2008, Morris et al. 2013, Cloonan et al. 2014). These simulators however are specific to a certain anatomical model or disease state. As previously discussed, access to experts to perform evaluation is generally limited, particularly in the early stages of medical device development. Current physiologically accurate simulators are usually laboratory based as emphasis is placed on the experimental integrity (such as for migration testing of stent grafts). This emphasis is often misguided particularly at the early stage of device development and when using early functional prototypes. One of the key advantages of the current work is the ability to easily configure the parameters for multiple anatomical models. For example, if a cerebrovascular model is required for device testing of aneurysm coils or clips, the perfusion rate to the brain is approximately 750 ml/min (Langlois et al. 2006). The anatomical model can be installed in the simulator with ease sue to modifications discussed in section The appropriate perfusion rate can then be set using a single mouse click on the LabView VI. The use of pulsatile anatomical simulators in the medical device industry for evaluation, training and demonstration is well known anecdotally. Little of this work is described in the literature due to the confidential nature of medical device companies. These simulators are either stationary or rigid anatomical models without flow that can be transported with ease.the portability of the current work, coupled with correct physiological flow and parameters is an improvement on contemporary designs. 112

127 Objective five: Simulated imaging and monitoring Research objective five identified the need to move away from direct visualisation of devices within hollow silicone models. The suspension of disbelief is a core element of any valid simulation. Direct visualisation impacts the operator in a number of ways that are detrimental to a suspension of disbelief. In live endovascular procedure the operators vision is fixed on the imaging screen, usually directly in front of the operator and at head height. If directly viewing a device in a silicone model, the operator will be looking downwards and to the side. This alters the normal posture and may induce a physical strain not present in clinical practice, negatively impacting the validity of real world use. The use of real time haemodynamic monitoring is novel in the current work. Experimentally focused pulsatile flow rigs (such as that used by the MSSI in Figure 8) do have pressure, flow and temperature sensors within the setup. However these are often recorded as raw data (voltage or current) and analysed offline after the experiment has concluded. The LabView VI created for this work (Figure 46) is presented in a similar style to a traditional cath-lab monitoring system such as the Picasso system pictured in Figure 56. One obvious limitation of the current work is the quality of simulated imaging presented in the literature compared to the imaging system used here. Figure 57 above shows the results of using plain film x-ray as an underlay for silicone anatomical models (Chong et al. 1998). This method, while producing impressive results, is difficult to reproduce now due to the move away from physical x-rays to digital capture methods. Underlay FIGURE 56: PICASSO HAEMODYNAMIC MONITOR 113

128 FIGURE 57: SIMULATED IMAGE USING X-RAY FILM AS BACKGROUND (CHONG ET AL. 1998) methods are also complicated by the need to match the refractive index of the circulating fluid to the material used in the anatomical model. As the use of 3D printed compound anatomical models is limited, research is required to qualify the refractive indexes of these new materials for this purpose. A further limitation of the simulated imaging system presented here is that it continuously presented live. In a clinical settings the fluoroscope is only activate briefly throughout the duration of the procedure to limit the exposure of the staff and patient to ionising radiation. The imaging is activated by a dual pedal footswitch, operated by the primary operator. The fluoroscope can be used in low power scouting type runs, or more detailed, higher power cine runs. All of these runs are recorded and stored on the display system and can be replayed and overlaid on screen as required. Future development of the current work should include the addition of a foot pedal to start/stop the fluoroscopic image. This requires the video image to be processed by a computer and code developed to work in conjunction with the foot controls and display, which was outside the scope of the current work. 114

129 Contribution of research The ultimate aim of applying human factors in the healthcare setting is to improve patient safety. Usability testing is the practical application of human factors in ensuring that medical devices reach the market that are both safe to the patient, and as easy to use as possible for the operator. The framework for integrating human factors into the design process (page 32) aims to formally illustrate the relationship, and importance, of usability testing in medical device design. The 3D printed compound anatomical models and simulator presented here will improve usability testing at all stages of the design process, particularly in the earlier stages. The deficiencies of using animal models or laboratory based simulators were previously discussed. The current work will provide a more realistic, higher fidelity method for early stage usability testing of endovascular medical devices. 115

130 6.2. Limitations Within the research presented here, several limitation were identified: The quality of the compound anatomical models is directly related to the quality of the medical image provided. Higher resolution scans with smaller pixel size and thinner slice thickness will improve the resulting compound models. The mechanical material properties for calcifications and thrombus are not well characterised. For this work, the softest material was used for thrombus and the hardest rigid material used for calcifications. In reality, the composition of both plaques and thrombus are not homogenous. Further research is required to characterise these materials against those generated on the Connex system. 3D printed compound anatomical models are more difficult to clean than those created with the lost wax method, as the support material must be removed by hand. An external laptop is currently required to run LabView; in future iterations an on-board computer is desirable to make it a truly standalone unit. The CCD camera is rigidly mounted in the current simulator; ideally the camera could be mechanised so that the operator can vary the view as with real C-Arm systems. 116

131 6.3. Future work In light of the work presented here, the following future work is proposed: In order to validate the compound anatomical models for both dimensional accuracy and structural integrity, it is proposed to complete a CT scan on a 3D printed AAA model using the same parameters as the patient scan that the model was generated from. Development of a higher fidelity simulated imaging system is required to bring the current simulator in line with contemporary work. Characterisation of thrombus and calcifications in comparison to materials created by the Connex system is required for future validation of compound anatomical models. A need exists for more complete anatomical phantoms for surgical and radiological training. 3D multi material printing offers a mechanism to produce these large scale anatomical phantoms with relative ease. 117

132 Chapter 7: Conclusions In this thesis, methods to segment and produce 3D printed compound anatomical models were presented. The construction of a pulsatile usability test bed is also described. The major conclusions are summarised as follows: Compound anatomical models can be digitally segmented from DICOM imaging using the Mimics software program. Compound anatomical models with multiple anatomical structures requiring different material properties can be created using a Polyjet 3D printing system. Currently, the Object Connex 3D printer in the most suited to the production of compound anatomical models. Real time haemodynamic and simulated fluoroscopic imaging systems can be integrated into a portable usability test bed. 118

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148 Appendix I: Arduino Code const int steppin = 13; const int dirpin = 12; const int limswitch = 11; const int bpm = A0; const int volume = A1; const int pace = 7; void setup(){ Serial.begin(9600); pinmode(limswitch, INPUT); pinmode(steppin, OUTPUT); pinmode(dirpin, OUTPUT); pinmode(bpm, INPUT); pinmode(volume, INPUT); pinmode(pace, INPUT); delay(500); digitalwrite(dirpin, LOW); // Output Stroke for (int i=0; i<steps; i++) { digitalwrite(steppin, HIGH); delaymicroseconds(5); digitalwrite(steppin, LOW); delay(2); } delay(20); } int limval = digitalread(limswitch); while (limval == LOW){ digitalwrite(dirpin, LOW); digitalwrite(steppin, HIGH); delaymicroseconds(5); digitalwrite(steppin, LOW); delaymicroseconds(3); delay(15); limval = digitalread(limswitch); } } void loop(){ int vol = analogread(volume); vol = map(vol, 0, 1023, 0, 80); Serial.print("ML \t"); Serial.println(vol); int long length = (float)((vol*1000l)/(3.14* )); Serial.print("Stroke \t"); Serial.println(length); int STEPS = (float)(length/0.1543); Serial.print("Steps \t"); Serial.println(STEPS); int Beats = analogread(bpm); Beats = map(beats, 0, 1023, 1, 120); Serial.print("BPM \t"); Serial.println(Beats); int long LabVal = (float)((60000l/beats)- ((STEPS*4.010)-(20*Beats))); Serial.print("Delay \t"); Serial.println(LabVal); Serial.println("************"); digitalwrite(dirpin, HIGH); // Input Stroke for (int i=0; i<steps; i++) { digitalwrite(steppin, HIGH); delaymicroseconds(5); digitalwrite(steppin, LOW); delay(2); } xvi

149 Appendix II: LabView Code xvii