UNIVERSITÀ DEGLI STUDI DI PISA
SCUOLA DI INGEGNERIA
Corso di Laurea in Ingegneria Biomedica
Dipartimento Ingegneria dell’Informazione
Tesi di Laurea Magistrale
I
MAGE
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ASED
M
ECHANICAL
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HARACTERIZATION OF
L
ARGE
B
LOOD
V
ESSELS
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A
C
OMBINED
A
PPROACH OF
E
XPERIMENTAL AND
C
OMPUTATIONAL
T
OOLS
RELATORI: CANDIDATO:
Ing. Vincenzo Positano
Benigno Marco Fanni
Ing. Claudio Capelli
Ing. Simona Celi
To my family, my love, my comrades.
I
NTRODUCTIONEngineering models can support innovations and improvements in the field of medicine. Over the last twenty years, the cardiovascular field has seen a flourishing of advances which included the design of more effective and less invasive devices. In this context, percutaneous pulmonary valve implantation (PPVI) was introduced to handle several congenital heart valve pathologies by means of a minimally invasive approach. Using a valve-equipped stent, the replacement of the dysfunctional valve is now possible avoiding open-heart surgery and this allow a partially recover the physiological heart function of the patient. In order to improve the efficacy of this or similar percutaneous interventions, patient-specific simulations could be used to predict the outcome of the procedure and the performance of the device once in contact with the vessel wall of the patient.
The planning of such intervention is currently based on imaging techniques such as magnetic resonance. Such imaging techniques provide information on the anatomy and on the function. While the relatively high spatial resolution generally allows fully understanding of the morphological characteristics, it is difficult however inferring data on the mechanical characteristics of the implantation site. Hence, this lack of precise information on the material properties is still the critical point for the validity of computational simulations. In fact, material parameters are usually taken from literature for FE simulations, while patient-specific geometry is provided by imaging. A strong validated image-based framework is still lacking. In this thesis, I investigated a combined approach of experimental and computational tools to infer mechanical properties of a model of a vessel by means of Magnetic Resonance (MR) imaging, in order to put basis for an image-based framework able to recognize the patient-specific material properties of large blood vessels, which would strongly enhance clinical decisions for percutaneous treatments.
M
ATERIALS ANDM
ETHODSAn experimental mock circulatory system (MCS) has been setup to test 3D-printed distensible models of blood vessels under cardiac pulsatile conditions. The
material used for the phantoms was TangoPlusBlack FLX980 (TangoPlus), printed by using PolyJet technology with two geometries: a hollow cylinder (15 mm length, 12.7 mm internal diameter and 2 mm thickness) and a patient-specific pathological pulmonary artery.
The MCS was powered by a pulsatile pump (Harvard apparatus pulsatile blood pump, Harvard Apparatus, USA) while flow and pressure information were measured by a flowmeter (Transonic System Inc., USA) and a catheter pressure (Opsens Inc., Canada), conveniently calibrated before their usage, both positioned in proximity of the phantom. After preliminary MCS tests, the circuit was positioned in the MR scanner (Siemens Avanto 1.5 T, Siemens AG, Germany) room in order to acquire Phase Contrast (PC) MR images of the phantom, firstly of the cylindrical sample and then of the patient-specific one, while sensors monitored and registered flows and pressures data by using a Biopac MP150 (Biopac System Inc., USA).
Cross-sectional images were acquired in through plane modality in the middle of the cylinder, while three different sections were acquired for the patient-specific phantom (proximal, stenotic and distal sections). PC MR technique was chosen in order to obtain the dynamic curves of flows and areas during the cardiac cycle imposed by the pump. The acquisition of each phantom was conducted in two different flow conditions by tuning the settings of the pump.
In order to get a first valuation of material properties of TangoPlus in terms of Young's modulus, the flow-area (QA) loop method, commonly applied on ultrasound (US) images, was used by analyzing the post-processing results of MR imaging data.QA loop method is based on the evaluation of the pulse wave velocity (PWV), from which the Young's modulus E is finally computed. The elastic modulus of TangoPlus derived from QA loop method was compared with the results of uniaxial tensile tests.
For this purpose, fifteen TangoPlus dogbone-like specimens were 3D-printed with different material fibre orientation (in order to assess the contribution of potential anisotropies deriving from printing direction). The specimens' set was composed by five samples with fibre orientation along x-axis, five along y-axis and five along 45° direction. In order to conduct the tensile tests, an extensometer system has been set up, composed by the testing machine (zwicki-Line Materials Testing Machine, Zwick/Roell, Germany) to pull the samples, a camera (HD Webcam C525, Logitech, Sweden) for tracking the markers attached on the sample in order to evaluate the strain
a laptop for post-processing.
Finally, Finite Element (FE) simulations were conducted in Abaqus (Dassault Systèmes, France) to replicate the conditions of the mock loop experiment by means of both structural and fluid-structural interaction (FSI) analyses. Structural simulations were run on both cylindrical and patient-specific geometries by applying on their inner surfaces the pressure loads measured by the catheter pressure during the MR imaging acquisition of the phantoms.
FSI simulations were conducted only on the cylinder by coupling the flow information registered by the flowmeter during the mock loop experiments. The influence of different material properties was explored taking into account data from literature, results from QA loop method and tensile tests.
R
ESULTSResults from MR imaging acquisition were post processed to calculate areas and flows dynamic curves as extrapolated from the segmentation of the images (Segment, Medviso, Sweden). Data of the imaging post-processing were used as input for the QA loop method, which computed the PWV as the slope of the linear part of the flow versus area scatter plot. From the obtained PWV value, the Young's modulus of the TangoPlus was estimated to be 0.221±0.041MPa.
The tensile tests conducted on the specimens resulted in an elastic modulus equal to 0.5001±0.0173 MPa. In addition, no significant anisotropy derived from the 3D printing direction was found.
The 0.5 and 0.22 MPa Young's moduli were assigned as material parameters of two linear elastic models for the FE simulations. Results were compared to estimate the relative error between the maximum area obtained from imaging segmentation and that resulted from structural simulations, evaluated in the middle section of the cylinder and in the stenosis of the patient-specific model. For the cylinder, the relative error measured 20.11% for condition 1 and 33.72% for condition 2. For the patient-specific phantoms, errors found to be 1.13% and 31.33%, respectively for conditions 1 and 2.
Results from FSI simulations consisted in the relative errors for both areas and flows but evaluated only for the cylindrical model. For 0.5 MPa model, errors found to be 2.16% for the area and 24.71% for the flow for condition 1, while 8.44% and 29.71%
respectively for condition 2. For the 0.22 MPa the errors were higher and for condition 1 measured 9.55% and 29.71% (area and flow), while for condition 2 I obtained 13.92% and 35.60%.
C
ONCLUSIONSIn this work, an experimental set-up has been developed in order to test a selected material for the inferring of its mechanical properties by means of MR imaging. TangoPlus phantoms were 3D-printed and inserted in a mock loop, then images were acquired in terms of PC MR imaging. Flow-area loop method was applied to the segmentation results of the imaging in order to evaluate the Young's modulus. As verification of the flow-area loop analysis result, uniaxial tensile tests were conducted on fifteen TangoPlus samples. Finally, FE simulation were run by assigning the material properties from tensile tests and flow-area loop method in order to replicate the areas and flows variations obtained from imaging post-processing.
Structural simulations' outcomes showed to be less accurate, as expected. FSI simulations were closer to replicate the area and flow measures obtained by the imaging data segmentation, especially for the areas, where the relative error was lower. The Young's modulus computed by the flow-area loop analysis resulted to be underestimated. Further investigations should be conducted in order to assess the reasons of such differences. In particular, a further evaluation of this method should be carried out comparing its use on US images, on which the flow-area loop analysis is validated, and MR imaging data. Furthermore, others rubber-like materials could be tested, following the same flowchart of this thesis, with the aim of developing an image-based framework able to characterize the mechanical behaviour of a patient-specific vessel.
Succeeding in developing a framework like this would improve advances in both device design and clinical decisions towards personalised care solutions with new modelling environments for predictive, individualised healthcare to guarantee better patient safety and efficacy.
ABSTRACT
LIST OF CONTENTS
LIST OF ABBREVIATIONS
CHAPTER 1 - INTRODUCTION ...9
1.1 Introduction ...10
1.2 Aims of the thesis ...12
1.3 Outlines of the thesis ...13
CHAPTER 2 - THE CLINICAL PROBLEM ...15
2.1 Introduction ...16
2.2 Congenital hearts diseases ...17
2.2.1 Incidence and consequences ...17
2.2.2 Tetralogy of Fallot ...17
2.2.3 Coarctation of the aorta ...18
2.2.4 Anomalous pulmonary venous return ...19
2.3 Percutaneous pulmonary valve implantation ...20
2.3.1 Introduction ...20
2.3.2 From surgical to percutaneous ...21
2.3.3 Devices for PPVI ...22
2.3.4 Procedure description ...24
2.4 Patient-specific FE models as clinical support ...26
2.4.1 Introduction ...26
2.4.2 FE methods for percutaneous valve implantation ...27
CHAPTER 3 - MATERIALS AND METHODS ...30
3.1 Introduction ...31
3.2 The mock circulatory loop ...32
3.2.1 Background ...32
3.2.2 Mock circulatory system description ...33
3.3.1 Calibration of the flowmeter ...37
3.3.2 Calibration of the catheter pressure ...38
3.4 Phantoms' fabrication ...40
3.4.1 3D printing ...40
3.4.2 Cylindrical phantoms ...41
3.4.3 Patient-specific phantom ...42
3.5 Magnetic Resonance Imaging ...46
3.5.1 Physical principles ...46
3.5.2 Image formation ...48
3.5.3 Phase contrast images: overview ...49
3.6 Phantoms' imaging acquisition ...52
3.6.1 Mock loop positioning ...52
3.6.2 Cylindrical phantom's acquisition ...53
3.6.3 Patient-specific phantom's acquisition ...56
3.7 The flow-area loop method ...59
3.7.1 Introduction ...59
3.7.2 Pulse wave velocity ...59
3.7.3 Image-based methods for PWV valuation ...60
3.7.4 From PWV to Young's modulus ...61
3.8 Uniaxial tensile tests ...63
3.8.1 Introduction ...63
3.8.2 Tensile testing ...63
3.8.3 Samples' fabrication ...65
3.8.4 Extensometer system set-up ...67
3.9 Finite element simulations ...71
3.9.1 Introduction ...71
3.9.2 Materials ...72
3.9.3 Finite element structural simulations ...75
3.9.4 Finite element FSI simulations ...80
CHAPTER 4 - RESULTS ...82
4.1 Introduction ...83
4.2 Imaging segmentation ...84
4.2.4 Area smoothing ...89
4.2.5 Results from cylindrical phantom ...91
4.2.6 Results from patient-specific phantom ...93
4.3 Flow-area loop analysis ...96
4.4 Results from tensile testing ...98
4.5 FE simulations' results ...101
4.5.1 Structural simulations under physiological pressure ...101
4.5.2 Structural simulations under experimental pressures ...102
4.5.3 FSI simulations ...104
CHAPTER 5 - DISCUSSION AND CONCLUSIONS ...108
5.1 Introduction ...109
5.2 Discussion ...110
5.2.1 Mock circulatory system ...110
5.2.2 MR imaging segmentation...110
5.2.3 Flow-area loop method ...111
5.2.4 Uniaxial tensile tests ...112
5.2.5 Finite element simulations ...112
5.3 Conclusions and perspectives ...115
BIBLIOGRAPHY ...116
L
IST OFA
BBREVIATIONSAR Aortic Root
AW Aortic Wall
CHD Congenital Hearts Disease DC Distensibility Coefficient
FA Flip Angle
FDM Fused Deposition Modeling
FE Finite Element
LVAD Left Ventricular Assist Device MCS Mock Circulatory System
MR Magnetic Resonance
PA Pulmonary Artery
PAPVR Partial Anomalous Pulmonary Venous Return
PC Phase Contrast
PLA Poly Lactic Acid
PPVI Percutaneous Pulmonary Valve Implantation PWV Pulse Wave Velocity
QA Flow-Area
ROI Region Of Interest
RV Right Ventricle
RVOT Right Ventricular Outflow Tract
TAPVR Total Anomalous Pulmonary Venous Return TOF Tetralogy of Fallot
US Ultrasound
VENC Velocity Encoding
CHAPTER 1
CHAPTER 1 - INTRODUCTION
10
1.1 INTRODUCTION
Biomedical engineering is the discipline which applies the concept and methods of physical sciences and mathematics to problems related to the medical field [1]. Biomedical engineering is by definition a multidisciplinary subject, which can run from mechanics to informatics, involving electronics, signal processing, material science and so forth. Its main purpose is to improve the healthcare, from the diagnostic point of view to therapeutic one.
This thesis focuses on the therapeutic field and, in particular, aims to explore the potential of computational analysis together with advanced imaging techniques, which are able to generate very accurate 3D anatomical patient-specific models. Ultimately, computational analysis can be used to plan personalized intervention.
Finite Element (FE) method is a technique of computational modelling, which allows complex problem to be solved by dividing continuous systems into a discrete number of components. Numerical models have been defined as the third pillar of science and engineering, achieving the fulfilment of the two more traditional disciplines, which are theoretical analysis and experimentation [2]. Advances in computational techniques have already improved the understanding of the biomechanical behaviour of healthy and diseased cardiovascular system. These models give the possibility to quantify physiological responses of the cardiac system under normal, diseased, and surgically-altered states, to improve vascular disease diagnosis, surgical planning, and prosthesis design, to assist the procedural phase, and even to facilitate the development of personalised treatments.
In this work, computational models have been set up to study an approach to infer mechanical properties of blood vessels by post-processing of Magnetic Resonance (MR) images, with the aim to improve effectiveness of minimally invasive cardiovascular interventions.
From the clinical point of view, the field of percutaneous valve implantation, a non-surgical minimally invasive technique of heart valve replacement was analysed, in particular the percutaneous pulmonary valve implantation (PPVI).
In general, this technique is based on the concept that a heart valve, the pulmonary one for PPVI, sewn inside a stent, can be reduced in size by crimping it onto a catheter,
11
and then introduced through a peripheral vessel to the desired implantation site in the heart. Inflation of the balloon, or release of a constriction sheath, deploys and anchors the valve-equipped stent within the old dysfunctional valve. This simple principle has formed the basis for successful clinical programmes, which have markedly changed the approach to treating cardiac valve dysfunction. The success of this technique is strongly based on the careful selection of the patient and the assessment of the mechanical characteristics of the implantation site for the device. In fact, the PPVI device works properly only if the main pulmonary artery has adequate level of distensibility. The quantification of the vessels’ distensibility via non-invasive technique can therefore improve the current selection of patients for PPVI.
Together with the computational modelling, this thesis has involved a broad range of different techniques and methodologies: the set-up of a mock circulatory loop, involving the 3D printing of rubber-like phantoms; Phase Contrast (PC) imaging; post-processing of the acquired imaging stack, in terms of segmentation and analysis of area-flow loop; uniaxial tensile tests of specific dogbone-like specimens for the mechanical properties determination of the rubber-like material used for phantoms 3D printing.
CHAPTER 1 - INTRODUCTION
12
1.2 AIMS OF THE THESIS
The principal aim of this work is inferring mechanical properties of blood vessels using post-processing analysis of PC MR images. Coupling the characterization of material properties derived by free radiation MR technique and the accuracy of patient-specific 3D geometry models, it could be possible modelling the interaction of a certain device into the vessel of that specific patient. From a clinical point of view, this would be extremely useful during the planning phase of minimally invasive percutaneous intervention.
The specific objectives are:
To develop a mock circulatory loop system using 3D-printed rubber-like phantoms;
To acquire morphological and PC MR images of phantoms;
To post-process the imaging stack and analysis of flow and area dynamic curves;
To test rubber-like phantoms manufacturing techniques with uniaxial tensile tests;
To validate algorithm for inferring mechanical properties or arterial vessels against standard experiments;
13
1.3 OUTLINES OF THE THESIS
The thesis is developed into four main chapters, in which methods and results are presented before a dedicated chapter of Conclusions.
In the Chapter 2, I introduced the clinical problem which motivates this research, providing a background of the percutaneous treatment of pulmonary valve dysfunction from an engineering point of view, presenting also the possibility to improve the benefits of the patients.
Chapter 3 provides the description of materials and methods involved in this research. Firstly, I present the mock circulatory loop, of which I describe the signal components constituting the hydraulic circuit, starting from the pulsatile pump to the phantoms and measurement sensors. A short description of 3D printing technologies involved in the phantoms creations is also presented, together with the methodologies used to calibrate the flow and pressure sensors. Then, the MR imaging acquisitions phase is presented. After providing a brief description of the physical principles of this imaging technique, in particular PC sequences, I present the acquisition of the mock loop system inside the scanner room. Afterward, I present the arrangement of specific 3D-printed specimens, made by the same rubber-like material used for the phantoms, for uniaxial tensile tests. Finally, the last section of this chapter focuses on the FE methods, starting from the material models used in literature to simulate blood vessels. Then, I present the computational model of the phantoms inserted into the mock loop, followed by the structural, fluid dynamics and fluid-structural interaction simulation planning.
Chapter 4 includes all the results of the thesis: imaging stack post-processing, analysis of flow-area loop, to infer a first evaluation of rubber-like material stiffness; results of tensile tests for a second value of stiffness; FE simulations results varying the material properties of computational model, assigned according to literature and mechanical values from tensile tests and flow-area loop analysis.
The next scheme shows a recap of the structure of the thesis [Fig. 1.1], which starts whit the phantom 3D-printing, continues with the mock-loop set up and MR imaging acquisition, goes on with post-processing imaging analysis and tensile test, and
CHAPTER 1 - INTRODUCTION
14
finishes with FE simulations, reaching the results and conclusions, in which the coherence of the outcomes are discussed.
Finally the Chapter 5 discuss the main findings of this thesis with ideas about future perspectives.
CHAPTER 2
CHAPTER 2 - THE CLINICAL PROBLEM
16
2.1 INTRODUCTION
In this Chapter, I introduce the clinical background which motivated this research. Hence, I start describing briefly the congenital heart diseases, their incidence and consequences. In particular, I focus on the principle diseases, including the Tetralogy of Fallot (TOF), which is the pathology for which percutaneous pulmonary valve implantation (PPVI) was developed.
In the third Section, I focus on the PPVI itself, outlining procedure and devices available to date, highlighting the main drawbacks which still limit its use on a large number of cases.
In the fourth and last Section, I provide a description of how computational methods can be used to support clinical decision. In particular, I focus on the integration of advanced imaging data which allow acquisition of high resolution 3D geometries, which are essential for setting up patient-specific simulations. Finally, I highlight the current main gap that this thesis tries to address, namely the challenges to translate into a model the mechanical characteristics of each patient vessel.
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2.2 CONGENITAL HEART DISEASES
2.2.1 I
NCIDENCE AND
C
ONSEQUENCES
Congenital heart diseases (CHD) are the most common type of birth defects, affecting about 1% of newborns per year [3]. CHD are anatomic lesions characterized by abnormal size, bad connections or communications of the chambers of the heart or the arteries and veins adjacent to the heart. The physiologic consequences of CHD are several, going from asymptomatic murmur detected only incidentally in adulthood, to severe cyanosis requiring urgent surgical intervention in the neonatal period. In particular, CHD can be divided into three representative specific diseases, presented in the following.
2.2.2 T
ETRALOGY OF
F
ALLOT
TOF is one of the most prevalent CHD requiring intervention in childhood. It usually consists of four defects. The primary anomaly is the abnormal formation of the infundibular septum, which is a muscular wall between right and left ventricles. This abnormality causes obstruction of the right ventricular outflow tract (RVOT), and as consequence the blood cannot correctly flow into the pulmonary artery (PA). Then, there is the so-called ventricular septal defect, which is given by an abnormal communication between the ventricles, causing an enlargement of the aorta and its overriding of the ventricular septum. Finally, the right ventricle muscle bulk appears increased, termed right ventricular hypertrophy. Figure 2.1 shows normal versus pathologic heart in TOF.
CHAPTER 2 - THE CLINICAL PROBLEM
18
The mortality rate for untreated cases is about 50% by 6 years of age, even if with the introduction of modern surgical techniques a 40% reduction in death association with TOF was registered [4] and now patients survive well into adulthood [5].
However, patients with repaired TOF need clinical and imaging follow-up to avoid the usual postsurgical complication, of which the most common are pulmonary regurgitation and PA stenosis.
2.2.3 C
OARCTATION OF THE
A
ORTA
Coarctation of the aorta, also called aortic narrowing, is a condition whereby the aorta is narrow [Fig. 2.2], usually in the area where the ductus arteriosus1 inserts. It is a fairly common congenital defect with an estimated incidence of 1 in 2500 newborns [6]. Aortic coarctation may occur alone, or be present in association with other CHD, such as bicuspid aortic valve.
Figure 2.2 - Representation of the coarctation of the aorta.
1
The ductus arteriosus is a blood vessel connecting the pulmonary artery to the proximal descending aorta.
19
The mortality rate approaches 80% by 50 years of age, due to aortic rupture or heart failure, but recent interventions with stenosis repair has resulted in significantly reduced patient morbidity [7]. However, constant follow-up is required in order to prevents complications, which include re-coarctation, aneurysm formation and aortic rupture.
2.2.4 A
NOMALOUS
P
ULMONARY
V
ENOUS
R
ETURN
Anomalous pulmonary venous return is a relative rare congenital anomaly which is divided into partial and total, according to the amount of venous return.In partial anomalous pulmonary venous return (PAPVR) the blood flow from some, but not all, pulmonary veins drain abnormally into a systemic veins or the right atrium. Then, blood oxygenated by passage through the lungs recirculates back through the lungs instead of supplying oxygen to the rest of the body, constituting a left to right shunt. Patients with PAVR are usually asymptomatic and are often diagnosed after childhood, even if some patients may come to attention prenatally or during infancy.
Total anomalous pulmonary venous return (TAPVR) is even rarer than the partial, where the totality of pulmonary venous drainage makes connections to the systemic venous circulation or right atrium [Fig. 2.3].
Figure 2.3 - Representation of total anomalous pulmonary venous return.
The management of TAPVR depends on the amount of drainage and the level of veins obstructions. Infants affected by obstructed TAPVR are commonly symptomatic and are surgically corrected early in life, with long-term postsurgical survival rates over 83% [8].
CHAPTER 2 - THE CLINICAL PROBLEM
20
2.3 PERCUTANEOUS PULMONARY
VALVE IMPLANTATION
2.3.1 I
NTRODUCTION
The pathologies described in the previous Paragraph require specific treatments, in particular referring on the RVOT disorders, whom reconstruction forms an integral part of surgical correction in a wide spectrum of CHD [9-10-11].
In general, disorder of cardiac valves may result in stenosis, when the valve fails to open normally, or regurgitation, in which the valve does not get closed correctly [12]. Prior to the beginning of this century, open heart surgery with valve replacement or repair was the only therapeutic alternative for the treatment of patient with symptomatic regurgitant valve lesions and for patients with calcific aortic stenosis [13].
Focusing in particular on RVOT dysfunction treatment, such operations are associated with long-term complications, which may be pulmonary stenosis, resulting in right ventricular (RV) hypertension and arrhythmias requiring graft valve2 replacement [15]. Therefore, depending on the type of valve in the pulmonary position, two to four surgeries are required to replace implanted devices in a lifetime of a child [15-16]. However, relatively recently valvular heart disease started to be treated also with less invasive interventions by using percutaneous transcatheter techniques. Percutaneous valve implantation is the development of a foldable heart valve that can be mounted on an expandable stent, delivered percutaneously through standard catheter-based techniques and implanted within a diseased valve location.
The first percutaneous transcatheter replacement of the pulmonic valve was introduced by Bonhoeffer et al. [17], which opened new perspectives by using the same technique also for other cardiac positions, such as aortic [18-19] and mitral valve locations [20]. Referring on pulmonary valve intervention, this technique is today usually known as PPVI.
Since its first successful attempt, several other interventions were carried on, demonstrating the feasibility of PPVI along the years [21-22-23], until our days. These
2
Interposition of graft valve between the RV and pulmonary artery has been one of the main surgical options for such defects.
21
successes can be expanded also to the other valvular heart districts, whereas all four valves are now managed with transcatheter methodologies [24].
2.3.2 F
ROM
S
URGICAL TO
P
ERCUTANEOUS
Obstruction of the RVOT is one of the anomalies of the Tetralogy of Fallot which can occur at valvular, sub- and/or supra-level, commonly leading to high right ventricular systolic pressure. In severe cases, this can cause RV hypertrophy and its progression can produce irreversible fibrotic change and RV failure. Pulmonary incompetence can generate RV volume overload, which can provoke RV dysfunction. This can lead to symptoms of reduce exercise tolerance, increased risk of ventricular tachy-arythmias and sudden death [25-26].
In this scenario, relief of the stenosis has to be performed in brief time. Surgical pulmonary valve replacement is now a low-risk intervention with a mortality of 1-4% and excellent mid-term survival [27-28], replacing the valve by using bioprosthetic
homograft3 and xenograft4, and mechanical valves. However, these devices present different disadvantage, such as limited availability for homograft and degeneration for xenograft, as well as their possible ruptures and other correlated problems.
The number of surgical replacement have decreased in favour of percutaneous implantation in the last years [29], such as the following plot shows [Fig. 2.4], and it is thought that the numbers of percutaneous interventions will increase even more when more devices will become available to the market.
Figure 2.4 - Percutaneous interventions versus surgical replacement during these years.
3
From homo, that means same, and graft, homograft are valve directly taken from human cadavers. 4 From xeno, different, and graft, xenograft are usually porcine valves.
CHAPTER 2 - THE CLINICAL PROBLEM
22
2.3.3 D
EVICES FOR
PPVI
Bonhoeffer's valve design was acquired by Medtronic and renamed it as Melody Valve (Medtronic Inc., Minneapolis, MN, USA) [Fig. 2.5]. The device is composed of a segment of bovine jugular vein with a native central valve that is sutured inside a balloon expandable stent CP stent5. Bovine jugular venous valves are available only up to 22 mm of diameter. Therefore, only RVOTs which are smaller than 22 mm of diameter can be treated with this percutaneous device. However, the valve's implantation was evaluated in further clinical reports demonstrating high procedural success and effectiveness in improving symptoms and eliminating pulmonary insufficiency[30-31-32].
Figure 2.5 - Melody Valve along long axis (left) and short axis (right)
To overcome some of the limitations related to the use of Melody Valve, new devices have been brought to the market. In particular, Edwards SAPIEN (Edwards Lifesciences INC., Irvine, CA) [Fig. 2.6] transcatheter heart valve was originally designed for the aortic position and it is now available also for the PA position.
Figure 2.6 - Edward SAPIEN valve along long axis (left) and short axis (right)
5
The CP Stent was designed by Dr. John Cheatham and it is made of a 0.013'' thick platinum/iridium wire mesh arranged in a zig-zag pattern. Every point of intersection is laser welded and coated with 24-carat gold.
23
The SAPIEN valve is a trileaflet bioprosthesis made of bovine pericardium that is mounted on a balloon-expandable stainless steel stent. The stent has a fabric cuff placed in the ventricular side that covers one half of the frame, limiting stent expansion and decreasing perivalvular insufficiency. The valve is available in 23 and 26 mm sizes allowing for implantation in larger RVOTs than the Melody valve.
In 2005, Garay et al. reported the first successful implantation of the SAPIEN valve in the pulmonic position of a 16-year-old patient with dysfunctional 24 mm homograft [33]. This was followed by the launch of the COMPASSION trial, a US multicenter study, to assess the safety and efficacy of the SAPIEN valve for the treatment of dysfunctional conduits demonstrating high procedural success and symptomatic improvement [34].
Today, these two devices are the only available on the market6, both suitable for only a limited number of patients requiring a new valve, but several others are under clinical trial [35], such as Venus P-valve (Venus MedTech, China) and Harmony (Medtronic, Ireland), both self-expandable devices [Fig. 2.7].
Figure 2.7 - Self-expandable devices: Harmony (left) and P-valve (right).
6
Even if the current generation NovaFlex+ catheter (Edwards Lifesciences Inc.) is utilized with the newer SAPIEN XT valve.
CHAPTER 2 - THE CLINICAL PROBLEM
24
2.3.4 P
ROCEDURE
D
ESCRIPTION
PPVI procedure consists in the placement of the valved stent within the existing degenerated valve or conduit via a catheter. A representative scheme of the procedure is displayed in Figure 2.8, where the valved stent-equipped catheter is positioned into the RVOT.
Figure 2.8 - Representative scheme of PPVI.
Transcatheter pulmonary valve replacement is usually performed under general anaesthesia with biplane fluoroscopic guidance. Femoral approach is usually chosen due to the setup of the catheterization lab; alternatively, jugular access can be preferentially selected to provide a more favourable anatomic curvature for device delivery. Right heart catheterization is initially performed to assess pressures and saturations with special attention to any relevant pulmonary branch stenosis. The balloon-tipped catheter is then replaced with a stiff guide wire and positioned into a distal branch pulmonary artery. Biplane RVOT angiography is performed to assess the proposed site for device implantation and quantification of pulmonary regurgitation. Implantation site dimension is then measured using a compliant sizing balloon. Pre-stenting with a bare-metal stent is routinely conducted in all cases of non-calcific outflow tract. The stent is typically deployed on a BiB (balloon-in-balloon) catheter (NuMED Inc., Hokinton, New York) to a diameter of up to 2 mm less than the original conduit size in stenotic conduits or slightly larger in conduits without stenosis.
The appropriate device size is then selected based on the diameter of the balloon at full inflation during pre-stenting. After appropriate device selection and preparation, the delivery system is advanced over the guide wire, bringing the valve into the
25
implantation site. Then, the BiB system is used with sequential inflation of the two balloons: first, partial deployment is achieved by inflation on the inner balloon is achieved by inflation of the inner balloon to confirm correct positioning; second, the outer balloon is inflated to complete deployment. Then, the balloons are deflated and the delivery system is carefully get out. Repeat angiography and pressure measurements are made to confirm a positive outcome [Fig. 2.9].
CHAPTER 2 - THE CLINICAL PROBLEM
26
2.4 PATIENT-SPECIFIC FE MODELS AS
CLINICAL SUPPORT
2.4.1 I
NTRODUCTION
Patient-specific modelling is an emerging field in biomedical engineering. The principal aim of this modelling is to improve patient healthcare by providing a personalised clinical treatment for each patient, starting from diagnosis and arriving to therapeutic field. Models of this kind are usually a combination of 3D reconstruction of a patient-specific anatomy, based on the modern available high resolution imaging data, coupled with numerical methods for solving the integrated mathematical problem of clinical treatment. In this context, FE methods has been used in life sciences to better understand the mechanics of biological tissues and their interaction with medical devices, such as those described previously7.
The approach of patient-specific modelling is multidisciplinary as it involves expertise in both medical and engineering subjects and requires a robust process, as visualized in the following workflow [Fig. 2.10].
Figure 2.10 - Workflow for patient-specific modelling [36]
27
2.4.2 FE
M
ETHODS FOR
P
ERCUTANEOUS
V
ALVE
I
MPLANTATION
PPVI is currently a low-risk and effective treatment for RVOT diseases [37-38]. The valves which are commercially available are limited just for those patients not presenting a large and dynamic RVOT, due to their dimensions, and device embolization remains the major problem in this kind of operations [37].
In this scenario, the development of more defined computational methods can contribute to study in details the mechanical interaction between device and implantation site. Up to date, computational simulations have shown the potentials to became effective tools to analyze biomedical devices [39] and their behaviour under patient-specific cardiovascular conditions [40-41-42-43-44].
In the field of percutaneous heart valves, computational models have been already implemented within patient-specific geometries to optimise device design [45] and evaluate risk of stent fracture [46]. However, patient-specific models for planning transcatheter valve implantation to date have not considered the individual implantation site material properties, which account not only for the vessel wall itself, but also for the surrounding support tissues. This lack of precise information can somehow limit the validity of the predictions of such computational tools.
At the moment, following the indications of the sizing balloon, the experience of the interventional cardiologist indicates whether to perform or not the PPVI, but a proper quantification of the mechanical response of the RVOT and surrounding tissue is still lacking. Notwithstanding all the advances in medical technology and engineering, techniques available to investigate the structure and the mechanical behaviour of human tissue are very limited in that they are highly invasive and consist mainly of histology and mechanical testing, for both aortic [47] and pulmonary tracts [48]. These techniques are not only invasive, since they need the tissue to be harvested, but they also represent the material property of a tissue ex vivo and when tissue physiological shape is altered, initial strain and residual stresses are removed, and physiological loading is absent. However, several material models have been developed to take into account the non-linear, hyperelastic and anisotropic behaviour of arteries [49-50-51-52].
Recently, few researches have been carried on in order to extrapolate the mechanical properties of vessel tissues by using indirect approaches. In 2012, Bosi et al. [53] have tried to infer mechanical behaviour of pulmonary valve implantation site by
CHAPTER 2 - THE CLINICAL PROBLEM
28
coupling systolic/diastolic dimensions and the pressure gradient in a linear elastic model and iterative tuning, assuming cylindrical thin-walled pressure vessel theory [Fig. 2.11].
Figure 2.11 - Laplace's law schematic. 𝑻𝜽 circumferential stress, P pressure, r radius and h wall thickness.
In 2013, Flamini et al. [54] proposed another non-invasive method to evaluate an approximation of the mechanical properties of aortic tissue, combining MR imaging and FE methods. In particular, a FE model was created from MR imaging data and loaded with systolic pressure. Then, the structural hyperelastic material properties were changed until average strains matched those measured by imaging [Fig. 2.12].
Figure 2.12 - Optimization routine workflow [2.52]
These two studies which describe methods to infer the mechanical properties of vessel tissue are representative of the effort to infer material properties of biological tissues in a non-invasive way. The majority of the studies reported in the literature adopts material constitutive law which describe generic vessels. These are usual approximated with mechanical models such as linear elastic, hyperelastic, either
29
isotropic or more complex anisotropic. In the Chapter 4, I took into account some of the most common material models and compared them against the results of the experiments.
However, a strong validated image-based method capable to infer the mechanical properties of vessel tissue, in order to simulate accurately eventual interactions between a certain device within a patient-specific geometry, is still lacking and it would be very useful from a clinical point of view. Such a method would be able not only to predict the feasibility of the intervention itself but also to offer the possibility to choose the better devices for this specific patient.
CHAPTER 3
31
3.1 INTRODUCTION
In this Chapter, I present the materials and methods used in this research. This study included both experimental and computational components which were adopted to infer the mechanical characteristics of vessels’ models in a non-invasive way.
In this context, first an MR compatible mock circulatory loop was set up in order to measure the distensibility of 3D-printed rubber-like phantoms. The idea was to build a simplified model of the cardiovascular system in which the attached phantom could expanded under the pressure generated by a pulsatile pump.
No control pressure was imposed, because the aim was just to measure motility and flow variations of the phantom under MR imaging during the cardiac cycle imposed by the pump, acquiring PC MR images.
The 3D printing technologies used to develop the phantoms and the others support pieces are described. In addition, specific dogbone-like specimens were 3D-printed, in order to do some uniaxial tensile tests for the rubber-like material involved. The extensometer system set up for these tests is discussed forward.
Finally, the set-up for FE simulations which mimicked the experiment are here presented: structural and fluid-structural interaction simulations.
In the following Sections of this Chapter, I here detail:
The configuration of the mock loop built up for this research;
The instrument systems used to measure flow and pressure variations imposed by the pulsatile pump, together with their calibrations;
The 3D-printed phantoms and support pieces fabrication, delving into the particular 3D printing technologies involved in this process;
A short overview of the MR imaging technique used in this study;
The phantoms acquisition phase set up,
The extensometer system used to perform uniaxial tensile tests;
The computational simulations developed by means of FE analysis, in particular structural and fluid-structure interaction (FSI) simulations.
CHAPTER 3 - MATERIALS AND METHODS
32
3.2 THE MOCK CIRCULATORY LOOP
3.2.1 B
ACKGROUND
Mock circulatory systems (MCS) are a valuable tool for hydrodynamic studies and in vitro device testing. The complexity of these systems can be adapted according to their use [55]. In the cardiovascular field, mock circuits have been used to test devices such as the intra aortic balloon pump [56], centrifugal and axial pumps [57], stents [58], and heart valves [59, 60].
In addition, the usage of MCS has been explored for conducting researches on a wide range of circulatory system-related pathologies. For example, in the left ventricular assist devices (LVAD) implantations scenario, a mock loop was used to evaluate intraventricular thrombus formation [61], or to correct right ventricular failure [62]. Even heat formation [63] or long-term durability [64] was investigated, such as potential mitral valve regurgitation [65] and a lot of other post-LVAD implantations. Examples of other pathologies of interest in this context are pulmonary arterial hypertension [66], unilateral stenosis in the pulmonary artery [67], unusual flow split following repair of transposition of the great arteries [68], and Eisenmenger’s syndrome [69].
Data gathered within a MCS can also serve as validation for computational studies [70]. These experimental setups can in fact be adapted to include phantoms attached to resistive and compliant elements according to a Windkessel model [55]. In this setting phantoms can be either an idealised test section [60, 71] or anatomical models which represent a specific patient [72, 73, 74]. Traditionally, rigid models have been manufactured using resins [74, 75] or glass [76]. While these are useful for visualisation studies, such as particle image velocimetry, they do not replicate the compliant nature of the vasculature and the associated Windkessel effect.
Manufacturing flexible models, nonetheless, can be challenging, and methods such as dipping [60, 77] or dripping [78] can be cumbersome, often with inadequate results. Different materials have been used for flexible vascular models, including silicone [70, 78, 79], polyurethane [80] and latex [81]. Recently, additive manufacturing techniques, commonly known as 3D printing, have allowed the fabrication of phantoms
33
which realistically mimic patients’ anatomies derived from medical images [82]. So-built phantoms can therefore be further investigated for validation studies [83], by mean of imaging techniques such as MR imaging. Materials used by 3D printers are plastics or rubber-like materials, and in any case MR compatible materials.
An example which is close to the anatomical target of this research, is the investigation of pulmonary hemodynamics using a compliant 3D-printed anatomical model [84], which involves both mock loop usage and MR imaging acquisition. In this scenario, the compliant attribute for 3D-printed material is important, in order to simulate the motility of real vessels under pulsatile pressure.
In this study we adapted a MCS to study the distensibility of 3D-printed rubber-like phantoms, of which the details are given in the following Paragraphs.
3.2.2 M
OCK
C
IRCULATORY
S
YSTEM
D
ESCRIPTION
In this study, the MCS was powered by a pulsatile pump mimicking the function of the heart, the Harvard apparatus pulsatile blood pump (Harvard Apparatus, USA).
The circuit in Figure 3.1 shows the structure of the simplified cardiovascular system with its four fundamental components: the pump, the piping system, the reservoir and the phantom to be tested. In addition, instruments apparatus allowed real time monitoring of the magnitudes of interests.
CHAPTER 3 - MATERIALS AND METHODS
34
Figure 3.2 - Harvard apparatus pulsatile blood pump.
The pump [Fig. 3.2] simulated the heart and his functions. In this system, Harvard apparatus allowed regulation of the flow by means of three different parameters: heart pump rate (bpm), stroke volume (cc) and the systolic-diastolic percentage phase ratio. In addition, it is also possible to monitor the output of the pump in terms of frequency. This has been useful during the MR imaging acquisition phase to trigger the scanner. The activity of the pulsatile pump is guaranteed by a system composed by a piston and two valves [Fig. 3.3], which replicate the functions of the cardiac valves. This system allows to simulate both systolic and diastolic phase, like in a real human heart system.
35
The fluid flowing in the circuit was chosen to be water. It is simple to get and to clean and it has similar behaviour under MR in comparison with blood.
In order to develop the closed loop of the human circulatory system, I used different kind of pipes.
In the scheme of Figure 3.1 different pipes are depicted with a colour code. The black ones are garden pipes which are attached to both the inlet and the outlet of the pump. These pipes had an internal diameter of 13 mm and 3 mm of thickness. The light blue pipes are smaller pipes, 2 mm of thickness and 9.5 mm of internal diameter. This size was compatible with the clamp-on flowmeter used in the experiment. The orange pipe is a compliant pipe which was used to prevent the negative pressure generated during the diastolic phase which could cause an undesired compression of the phantom.
The closed loop is so composed: a garden pipe connects the reservoir to the inlet of the pump. Another garden pipe origins from the outlet of the pump and it's connected to the small pipe which gets into the lumen of the bigger pipe. The small pipe connects to the phantom which connects to a second small pipe. This last one is connected to a normal garden pipe which bring the water to the reservoir. Then the water goes back to the pump.
The white box of the scheme is a reservoir, which took into account of the hydraulic human resistance. The height of the water into the reservoir was set at 7 cm in all the experiments, in order to have a fixed condition during the preparation part of the experiment. The reservoir was also useful to clean out air bubbles of the circuit, which would have affected the MR imaging acquisitions.
The core of the circuit (represented by a grey box in Figure 3.1) was a 3D-printed phantom made by TangoBlackPlus FLX 980 (TangoPlus), a particular rubber-like material, chosen for its capability to easily deform. The TangoPlus materials family was already investigated and found to be suitable for arterial phantoms [85], due to its distensibility properties. Two phantoms were tested in this project: a hollow cylinder (15 mm height, 12.7 mm internal diameter, 2 mm thickness) and a patient-specific geometry of a pathological pulmonary artery.
All the pipe connections were sealed to prevent leakages with PTFE1 (common Teflon) on the external surface of the small pipe and a cable tie shrinking the external surface of the bigger garden pipe. In addition, the connections between the phantom and
1
Polytetrafluoroethylene, a synthetic fluoropolymer of tetrafluoroethylene that has numerous applications.
CHAPTER 3 - MATERIALS AND METHODS
36
the small pipes were protected by a bandage on the phantom to avoid the tearing of the phantom due to the cable tie.
Finally, flow and pressure within the circuit were monitored by sensors (green boxes in Figure 3.1). Two sensors were used in this circuit: one flowmeter and one catheter pressure. The flowmeter kit was the Transonic 400-series multi-channel flowmeter consoles & modules for laboratory research (Transonic Systems Inc., USA). The catheter was the Opsens optical fibre catheter pressure (Opsens Inc., Canada) to measure the pressure. Both sensors were connected on the small pipe connecting the inlet of the phantom. In particular, the clamp-on flowmeter was connected in the proximity of the phantom, secured on the external surface of the small using an ultrasound gel to improve the signal coupling. The tip of the catheter pressure was inserted into the same small pipe through a pin-hole.
Their measured signals were acquired using a Biopac2 (Biopac System Inc., USA) connected via Ethernet to a laptop in which they were visualized in real-time through the Biopac AcqKnowledge software3.
2Biopac MP150 system. 3
37
3.3 CALIBRATION OF THE SENSORS
The calibration of the measurement system was performed prior to the experiment.
The flowmeter used in the experiment was a clamp-on one, made by Transonic. The equipment was composed by the sensor4 itself, the cable and the console (two modules5). The sensor is connected to the console, which is connected to the second channel of the Biopac. The mean value of the flow can be read from the module. The dynamic curve of the flow is get by the Biopac and displayed on a laptop using AcqKnowledge.
The catheter pressure used in the experiment was an optical fibre one made by Opsens within 9F6 catheter. The catheter is connected to the small Opsens console which is connected to the first channel of the Biopac.
3.3.1 C
ALIBRATION OF THE
F
LOWMETER
The Transonic flowmeter was calibrated with the "timed collection" method using a continuous pump made by Sicce7 (Sicce, USA), a measuring cylinder and a stopwatch. The pump has a variable flow setting and allows the users to have three different flow values (minimum, medium and maximum flow8). For the calibration, the pump was connected to pipe fitting with the flowmeter.
For each flow setting, the flow read from the console was compared to the flow calculated as ratio of the amount of water collected in the measuring cylinder and the time measured by the stopwatch. Each measure of each flow setting was repeated five times in order to ensure repeatability. The average of the measures of the flow fitted with the mean flow value read from the console for each flow setting.
The Biopac read voltage values and the voltage-to-flow ratio was derived by fitting the three points measured with a linear trend line. Results are shown in Figure 3.4.
4 Transonic ME 9PXL flow sensor. 5
Transonic TS410 Tubing Flowmeter Module.
6 French scale: 𝐹𝑟 = 𝐷 ∗ 3; e.g. 9F corresponds to 3 mm outer diameter catheter. 7 Sicce MI Mouse continuous pump.
CHAPTER 3 - MATERIALS AND METHODS
38
The function 𝑄 = 𝑓(𝑉) is resulted to be:
𝑄(𝑙/𝑚𝑖𝑛) = 3.56 ∗ 𝑉 − 0.03
Figure 3.4 - Flow (l/min) versus voltage (V) curve from calibration.
3.3.2 C
ALIBRATION OF THE
C
ATHETER
P
RESSURE
The calibration of the Opsens catheter pressure was made using the "column of water" method by knowing the relation:
10 𝑐𝑚𝐻20 = 7,355 𝑚𝑚𝐻𝑔
Hence, a hollow tube of 1.7 m was vertically positioned and filled with water. The catheter was inserted in a side branch on the bottom of the tube. The output of the sensor in voltage was acquired using AcqKnowledge. Calibration was performed by measuring pressure after variations of the water level from 152 cm to 12 cm at each 10 cm interval markings. The atmospheric pressure was also taken into account. The results of the calibration are shown in Figure 3.5. The pressure-to-voltage ratio is derived from this graph by fitting a linear trend line.
39
The relation 𝑃(𝑉) is found to be:
𝑃(𝑚𝑚𝐻𝑔) = 99.88 ∗ 𝑉 + 0.15
CHAPTER 3 - MATERIALS AND METHODS
40
3.4 PHANTOMS' FABRICATION
3.4.1 3D
P
RINTING
Both phantoms and connection pieces, described in the following paragraphs, used in this phase have been realized using 3D printing technology. In particular, the phantoms were printed using the machine Objet500 Connex (Stratasys, United States), which takes advantage of the PolyJet technology. The material used was TangoBlackPlus FLX890. The characteristics9 of TangoPlus are shown in Figure 3.6.
Figure 3.6 - Some TangoBlackPlus FLX980 characteristics.
PolyJet 3D printing works similarly to inkjet printing, but instead of jetting drops of ink onto paper, PolyJet printers jet layers of curable liquid photopolymer onto a build tray. The printing process is simple and it is composed by phase of pre-processing, production and support removal. During the pre-processing phase, Build-preparation software automatically calculates the placement of photopolymers and support material from a 3D CAD file. Then, in the production phase, the 3D printer jets and instantly UV-cures tiny droplets of liquid photopolymer. Fine layers accumulate on the build tray to create one or several precise 3D models or parts. Where overhangs or complex shapes require support, the 3D printer jets a removable support material. Finally there is the support removal, during which the user removes the support material manually, with water or in a solution bath. Models and parts are ready to be handled immediately after the 3D printing process, with no post-curing needed.
41
In addition to the phantoms in TangoPlus, it was necessary to manufacture two more pieces, in order to connect the patient-specific phantom to the mock loop. These two connections were 3D-printed by using another printer, the MakerBot Replicator Desktop 3D Printer (Stratasys GmbH | MakerBot Division, Germany), and the material was PLA (polylactic acid), a rigid material. This printer uses the fused deposition modeling (FDM), which creates its layers by melting thermoplastic material to a semi-liquid state and drawing out each layer before moving to the next layer. In order to do this, FDM uses two types of structures to complete a printing job: modeling material which finally becomes the finished product and support material which works as scaffolding. Once the model is out, the user can either separate the support material by breaking it off. Once the last step is complete, the 3D model is ready for use.
3.4.2 C
YLINDRICAL
P
HANTOMS
The experiment was conducted in different stages of increasing level of complexity. In the first phase of the experiment, the geometry chosen for the phantom was a simple cylinder. The shape was chosen to measure its distensibility in a controllable fashion. The geometry was designed and meshed using the Sketch Module of Abaqus10 (Dassault Systèmes, France). The same model described here was used for
the FE simulations discussed in the Section 3.9. The internal radius of the cylinder was chosen in order to fit with the rest of the mock loop and it was set to 6.35 mm. The length was set to 15 cm in order to manage the bounded edge effects and to acquire the variations of cross-sections by MR imaging acquisition phase, in the middle of the cylinder.
Initially, three different thicknesses were thought for the cylinder: 1, 2 and 3 mm. Once the part11 was designed in Abaqus, it was opportunely meshed and exported as
stl12 file. The three stl files with different thicknesses were sent to the Architecture Department of UCL to be 3D-printed using PolyJet technology. After preliminary testing, just the 2 mm thickness cylinder was taken into account for the experiment. It
10
Abaqus 6.14-5 version.
11 In the Abaqus jargon, the part is a component of the model. In this simple case, the model is composed by one single part, the cylinder.
12stl means STreoLitography, even if it's known also as Standard Triangle Language, and it is a file format widely used for 3D printing, such as in this case.
CHAPTER 3 - MATERIALS AND METHODS
42
provided a good compromise between safety and motility, without any breaking risk during the experiment and with a satisfying deformation level.
The 1 mm and 2 mm cylindrical phantoms, realized by using horizontal printing, are shown in Figure 3.7.
Figure 3.7 - Starting from the top: 1 and 2 mm cylindrical phantoms.
3.4.3 P
ATIENT
-S
PECIFIC
P
HANTOM
The second phase of the experiment consisted in testing a phantom which replicated anatomical characteristics of a real patient. In particular, I replicated a pathological pulmonary artery geometry obtained from a real patient.
The geometry, in form of stl file, was derived from the 3D volume reconstruction of high resolution Computed Tomography images of the patient. The model contains the RVOT, the main tract of pulmonary artery, except the valve, the stenosis and the first part of pulmonary branches. A thickness of 2 mm was assigned to the model using Meshmixer (Autodesk Inc., USA) as in the cylindrical phantom. The branches were clipped in order to simplify the connection of the patient-specific phantom to the rest of the circuit. Both geometries with thickness pre- and post-cut are shown in Figure 3.8.
43
The geometry was further modified to allow the connection of the patient-specific phantom to the rest of the circuit. Hence I chose to extend both the extremities with a cylindrical extrusion, using a specific tool available in VMTKLab (Orobix, Italy). Finally, these extrusions were further reinforced by increasing their thickness (i.e. 3 mm versus the 2 mm of the rest of the phantom).
The complete 3D-printed patient-specific phantom is shown below [Fig. 3.9].
Figure 3.9 - Final 3D-printed patient-specific phantom.
In order to connect the patient-specific anatomy to the rest of the circuit, it was also necessary to fabricate two extra components to adapt smoothly inflow and outflows. The next figures show the rendering of the model and the longitudinal sections equipped whit all the geometrical dimensions [Figg. 3.10-3.11].
CHAPTER 3 - MATERIALS AND METHODS
44
Figure 3.11 - Longitudinal sections showing the geometrical dimensions in mm of the connections.
Each piece has the extremity to be connected to the phantom characterized by a diameter dimensioned in order that the extremity could have been inserted into the cylindrical edges of the phantom leaving the smallest interstice possible between the surfaces. On the other side, the other extremity has to fit with the dimensions of the garden pipes. These piece were 3D-printed by using FDM technology and PLA as material [Fig. 3.12].
45
Both internal and external surfaces were covered with a coat of vinyl glue to make it water resistant.
These connections were accessible for the catheter pressure with a hole made in the inlet part in proximity of the connection with the phantom. The connection pieces and the final fully-equipped patient-specific phantom are illustrated in Figure 3.13.
CHAPTER 3 - MATERIALS AND METHODS
46
3.5 MAGNETIC RESONANCE IMAGING
3.5.1 P
HYSICAL
P
RINCIPLES
The imaging technique used in these experiments was MR imaging. This technique is based on the physic principle of the element resonance, in particular on the H+ (proton) resonance, of which the human body is rich.
Each element can be characterized by its gyromagnetic ratio, which is described as:
𝛾 = 𝑚 𝑝
In the above equation, 𝑚 represents the magnetic moment, because the particle is a moving electric charge, and 𝑝 is the angular moment, because at the same time is a rotating mass. In the human body, all the proton magnetic dipoles are randomly oriented in the space. The application of an enough strong13 static constant magnetic field 𝐵0 allows the nuclear spins to orient in the parallel14 direction of 𝐵0, which is aligned with
z axis in a Cartesian coordinate system. Such magnetic field is generated by the magnet
of the MR scanner. For the proton, the gyromagnetic moment is equal to 42.5 MHz/T. This quantity is important for the determination of another important parameter, the Larmor frequency, which is given by:
𝜔 = 𝛾𝐵0
It is also called precession frequency, because 𝜔 is the frequency at which the spin with a gyromagnetic moment 𝛾 and exposed to the magnetic field 𝐵0 rotates around its axis, with a precession motion. The Larmor frequency is in the band of radiofrequencies. For instance, having 𝛾 = 42.5 MHz/T and 𝐵0 = 1.5 T, we have a precession frequency 𝜔 = 63.87 MHz.
13 The intensity of the constant magnetic field goes about from 1.5 T to 7 T and depends on the application field.
14
Precisely, the dipoles get oriented in the parallel and unparallel direction, of which the parallel oriented dipoles are slightly more numerous, about 51% versus 49%.
47
If the spins exposed to the magnetic field 𝐵0 get excited by a magnetic pulse with radiofrequency 𝜔, the resonance phenomena is manifested. These magnetic pulses are generated by particular coils15, which can handle both transmitter and receiver works. The MR signal derives from the tissue response to these magnetic pulses. The temporal distance between a pulse and the next is called repetition time 𝑇𝑅, while the time been between the beginning of the pulse sent by the coil and the signal reception is named echo time 𝑇𝐸 [Fig. 3.14].
Figure 3.14 - Difference between TR and TE
Each proton can be described by its magnetization vector 𝑚, which is parallel to the magnetic field 𝐵0. The sum of all the vectors 𝑀 = 𝑚𝑖 is called net magnetization vector. Depending on the so called flip angle (FA), this vector get moved from the z direction. The FA is function of the duration pulse τ and the pulse frequency 𝑓 and it follows the linear law 𝐹𝐴 = 𝜏𝑓. Such as examples, after a certain magnetic pulse, 𝑀 can go laying in the x-y plane (𝐹𝐴 = 90°) or even getting completely inverted and aligned with -z direction (𝐹𝐴 = 180°).
After the end of the pulse, the net magnetization vector 𝑀 recovers his initial position, along direction z. The time 𝑇1 in which its z-component 𝑀𝑧 regains the initial amplitude and the time 𝑇2 in which the xy-component 𝑀𝑥𝑦 returns to 0 are features of the single tissues:
𝑀𝑧 = 𝑀0(1 − 𝑒−𝑡 𝑇1) 𝑀
𝑥𝑦 = 𝑀𝑥𝑦0𝑒 −𝑡
𝑇2
𝑇1 is named the spin-lattice relaxation time, because it depends on the interactions of the protons with the surrounding elements, while 𝑇2 is the spin-spin relaxation time
15
There are different kind of coils and the used type depends on the application: there are body coils, brain coils, single coils, phase coils, etcetera.
