O le k s a n d r B a ra n n y k 1 Mem. ASME Department of Mechanical Engineering, University of Victoria, P.0. Box 1700, STN CSC, Victoria, BC V8W2Y2, Canada
P e te r O s h k a i Mem. ASME Department of Mechanical Engineering, University of Victoria, P.0. Box 1700, STN CSC, Victoria, BC V8W2Y2, Canada
The Influence of the Aortic Root Geometry on Flow Characteristics of a Prosthetic Heart Valve In this paper, performance of aortic heart valve prosthesis in different geometries of the aortic root is investigated experimentally. The objective of this investigation is to estab lish a set of parameters, which are associated with abnormal flow patterns due to the flow through a prosthetic heart valve implanted in the patients that had certain types of valve diseases prior to the valve replacement. Specific valve diseases were classified into two clinical categories and were correlated with the corresponding changes in aortic root geometry while keeping the aortic base diameter fixed. These categories correspond to aortic valve stenosis and aortic valve insufficiency. The control case that corresponds to the aortic root of a patient without valve disease was used as a reference. Experiments were performed at test conditions corresponding to 70 heatslmin, 5 .5 Llmin target car diac output, and a mean aortic pressure of 100 mmHg. By varying the aortic root geome try, while keeping the diameter of the orifice constant, it w a s possible to investigate corresponding changes in the levels of Reynolds shear stress and establish the possibility of platelet activation and, as a result of that, the formation of blood clots. [DOT: 10.1115/1.4029747]
Introduction Prosthetic heart valves such as mechanical (MHV), polymeric (PV), and tissue valves have been used in heart valve replacement over many decades. Successful analysis of the flow through pros thetic heart valves depends on sufficient understanding of the con ditions under which natural valves function. Most disorders of the heart initiate within the left ventricle, as this chamber is subjected to the highest mechanical loads. The flow through the left ventri cle is regulated by the mitral and the aortic valves, which influ ence the inflow and the outflow conditions, respectively [1], The mitral and the aortic valves are the most commonly affected heart valves in a diseased heart, and they are responsible for 34% and 44% of morbidity [2,3], respectively. Much progress has been achieved in the development of artifi cial heart valves [4,5]. Nevertheless, serious problems still exist due to abnormally high-velocity gradients that are present within the jets that emanate from these prosthetic valves. These gradients result in elevated shear stresses that may cause red blood cell damage, platelet activation, and thrombus formation [6,7], which in turn lead to thromboembolism, hemorrhage, and tissue over growth. Flow-induced stresses in blood, acting on a cellular level, have been known to cause thrombus initiation within the compo nents of the mechanical valve prostheses [8], Regions of elevated stresses during the complex motion of the leaflets in some cases lead to structural failure of the MHVs. In vivo and in vitro experi mental studies have yielded valuable information on the relation ship between hemodynamic stresses and the problems associated with the implants [3,9]. When studying such complex phenomenon as blood flow through the heart valve, the choice of the methodology is critical. Two major approaches employed by researchers in the field of
‘Corresponding author. Manuscript received September 3, 2014; final manuscript received February 4, 2015; published online March 5, 2015. Assoc. Editor: Ender A. Finol.
Journal of Biomechanical Engineering
cardiovascular fluid mechanics are experiments and computa tional fluid dynamics (CFD) simulations. The experimental methodology had been used extensively in the investigation of various cardiovascular devices and novel intervention techniques. It is still a method of choice required for quantitative assessment of the hemolytic and thrombogenic poten tial when a novel cardiovascular device is undergoing a regulatory submission for a premarket approval [10]. Among several experi mental methods that allow flow visualization, digital particle image velocimetry (PIV) is particularly useful, as it can deliver global, quantitative flow images with high spatial and temporal resolution. In Refs. [11-14], PIV was used to quantify the velocity field in the vicinity of valve leaflets. The authors of Ref. [15] were among the first to apply PIV methodology to study flow through an artificial heart valve, and they outlined experimental challenges associated with it. The occurrence, severity, and distribution of viscous and Reynolds shear stresses (RSS) were successfully esti mated via PIV. Recently, microcomputed tomography and stereo scopic digital PIV techniques were used to investigate the hemodynamics of the mitral valve [16] and to evaluate a new design of bileaflet mitral valve prosthesis [17]. CFD methodology has been actively employed for the last dec ade to aid in the evaluation of medical devices and to improve safety and quality in the development of products and technolo gies. While still requiring extensive validation, CFD is a very powerful tool that can be successfully used in modeling of com plex, chemico-biological phenomena. The primary advantage of CFD over experimental fluid mechanics is its ability to provide fast analysis of the results of small changes in the design of medi cal devices. The authors in Ref. [18] used unsteady Reynoldsaveraged Navier-Stokes equations simulations in order to assess the flow through a 2D bileaflet MHV under steady inflow condi tions and estimate the influence of the flow field on the platelet activation mechanism. In Refs. [9] and [19], the CFD methodol ogy was applied to investigate complex flows through a bileaflet MHV in an axisymmetric aorta and within the total
Copyright ©2015 by ASME
MAY 2015, Vol. 137 / 051005-1
cavopulmonary connection, respectively. In Ref. [20], a full threedimensional (3D) fluid-structure interaction model of the aortic valve with coaptation and a compliant aorta was used to resolve the kinematics of the aortic valve and the details of leaflet coapta tion at the end of the closing phase of the cardiac cycle. The authors in Refs. [21] and [22] investigated the possible influence of different bifurcation stenting techniques on stent deformation as well as physical stress and drug elution via coupled mechanical-CFD numerical model. The primary consideration for the present study was the devel opment of anatomically correct benchtop experimental model with the ability to investigate the correlation between pathological changes in the geometry of aortic root induced by aortic valve ste nosis and insufficiency and presence of the abnormal flow patterns downstream of the recently implanted valve. The main design considerations for the experimental system included robustness, versatility, and compatibility of the model with the commercial left ventricular heart simulator. In addition to simulating anatomi cally correct geometries, the main physiological features of the healthy valve, such as nonnotensive pulsatile flow and pressure waveforms, were also implemented.
Experimental System and Techniques The experiments described in this paper were performed in the laboratory of ViVitro Labs, Inc., in Victoria. BC, Canada.
configurations while assessing their performance and function under simulated cardiac conditions. The pulse duplicator was powered by a piston-in-cylinder pump driven by a motor connected to the ViViTest data acquisition and control block, which controlled the pump. This setup allowed for various ventricular waveform states and beat rates while generat ing physiological pressures and flow regimes through the heart valve. The present investigation was focused on the hemodynamics of a prototype trileaflet PV, shown in Fig. 3(a). The valve, which had a nominal diameter of 19 mm, was mounted in a model of the aor tic root with severe pathological changes in the aortic root geometry. For each of the aortic root types, the flow fields were collected in two perpendicular data acquisition planes, as shown in Fig. 3(b). The start of the image capture was triggered by a syn chronization pulse from the pump controller. The PIV data were acquired at six representative phases of the cardiac cycle, indi cated by black circles in Fig. 4. These phases correspond to the opening acceleration phase (t/T = 0.05), the peak systole phase (Cl T = 0.13), the deceleration phase (tIT= 0.22), the closing phase (tl T = 0.29), and the leakage phase (t/T = 0.69). Here, t denotes time, and T = 860 ms is the period of the simulated cardiac cycle. Aortic Root Geometry. The aortic root is a part of the ascending aorta, shown in Fig. 5, with a combined length h = 80 mm [1,23,24], In this paper, the aortic root was defined by means of the following parameters [25,26]: D0 is the diameter of the orifice
Pulsatile Flow System. In Figs. 1 and 2, one can see the sche matic of the experimental apparatus that allows testing of the heart valves in the aortic configuration. The ViVitro pulse duplicator is a heart model that incorporates the functional capability to test two heart valve substitutes simultaneously in left or right heart
Fig. 3 (a) Orientation of the valve with respect to the left coro nary artery, the right coronary artery and the noncoronary cusp and the PIV data acquisition planes (dashed lines) and ( b ) sche matic of the prototype trileaflet PV Fig. 1 Schematic of the flow visualization setup (image cour tesy of ViVitro Labs, Inc.)
T im e (m s)
Fig. 2 Schematics of the test chamber and the aortic valve (image courtesy of ViVitro Labs, inc.) 051005-2
/ Vol. 137, MAY 2015
Fig. 4 Variation of flow rate as a function of time during a typi cal cardiac cycle. Black circles correspond to the phases of the cardiac cycle, at which PIV data were obtained.
Transactions of the ASME
Fig. 5 M ajor parts of the aorta and principal dim ensions of the aortic root sinuses
(also referred as aortic annulus diameter), DA is aortic diameter distal to sinus or diameter of the sinutubular junction, DB is the maximum projected sinus diameter, which approximate the diam eter of anatomic ventriculoarterial junction, LA is the length of the sinuses, and LB is the distance between D0 and DB. It should be noted that Da , the diameter of the orifice, was kept constant in order to interface properly with the ViVitro pulse duplicator. The primary objective of this study was to investigate the path ological changes in the dimensions of the aortic root due to aortic valve disease, such as valve stenosis and valve insufficiency, and to determine the influence of those changes on the appearance of abnormal flow patterns in the flow through the aortic valve which replaced the original, diseased valve. These pathological changes in extreme scenarios are often associated with patients morbidity due to aortic dissection, aortic incompetence caused by a dilated, aneurysmal aortic root, and severe aortic valve stenosis [27-30], In addition, deformation of the aortic root after valve replacement or structural dysfunction of the recently replaced bioprosthetic heart valve due to premature calcification associated with pure stenosis due to cusps stiffening is not uncommon [31,32]. Proper replication of the complex anatomy of the aortic valve and root is essential in many ways. First, the accurate measure ment of the aortic root geometry in clinical settings is required for evaluation of the patients with aortic valve or aortic root disease and determining whether the proposed device can be safely and successfully implanted, which is primarily related to the newly emerged transcatheter aortic valve implantation (TAVI) [25,27,33]. Second, the accurate representation of a complex aor tic root anatomy is essential in order to reproduce the internal physiological flow field correctly for any in vitro study. During just one cardiac cycle that includes the systole and the diastole phases, several complex flow events occur within the aortic root. The opening and the closing motions of the valve are supported through vortex formation within the aortic sinuses. In addition, when the valve is closed, previously formed sinus vortices are responsible for washout of the sinus cavities, so that the throm botic deposition does not occur. Among many techniques that are capable of a nonintrusive imaging of elements of the cardiovascular system, angiography is probably the most well known. The projectional radiography or angiography is a technique that employs a special dye in order to visualize the blood flow in arteries as well as the dimensions of the arteries. Nowadays, there are several highly sophisticated methodologies that are capable of accurate assessment of the major dimensions of the aortic root [25,27,34]. One of the most commonly used techniques for such assessment is transthoracic echocardiography (TTE). TTE can be applied to determine the
Journal of Biomechanical Engineering
severity of aortic valve/root disease and whether the patient is suitable for a TAVI procedure [27]. In addition to TTE, a contrast-enhanced multidetector computed tomography (MDCT), due to its accuracy, can be applied to determine the difference in aortic root geometry between patients with tricuspid and bicuspid aortic valve. Among the major drawbacks of MDCT is the exces sive radiation exposure to the patient and the risk of hypersensitiv ity to iodine contrast. As an alternative to MDCT, a 3D transesophageal echocardiography (TEE) has emerged and been demonstrated reliably to evaluate aortic root geometry [34]. The importance, of studying aortic root geometry associated with a particular clinical case in conjunction with a valve pros thetic in a single study, was stressed in Ref. [26], The authors showed a concise quantification of the aortic root geometry changes due to valvular diseases. Geometric parameters and alge braic quotients were obtained from 604 independent angiographic films collected from patients with various types and degrees of valvular diseases [26]. These calculated quantities were used in the present work in such way that the resulting aortic root geome try was compliant with the existing ViVitro pulse duplicator sys tem. By setting the orifice diameter to D0 = 24 mm and the wall thickness to A = 1 .5 mm, the dimensions listed in Table 1 were derived from relationships originally presented in Ref. [26]. In this paper, the geometries of the aortic root that corresponds to the aortic valve stenosis and aortic valve insufficiency [26] are referred as dilated aortic root and constricted aortic root, respectively. Based on the dimensions presented in Table 1, three types of geometries that represented aortic root and ascending aorta system were built through the process of stereolithography (SLA). DSM Somos® Watershed© XC 11122 (Watershed), an optically clear resin with acrylonitrile butadiene styrene-like properties, good tem perature resistance, water resistance, and high durability was used as a base material. This material is also a Class VI Approved Medi cal Grade material making it suitable for any medical applications. Watershed material has an index of refraction (IoR) of 1.51. Table 1 Principal dim ensions of the aortic root associated with clinical cases of heart disease
D0 (mm) Da (mm) DB (mm) LA (mm) LB (mm) Normal Dilated aortic root Constricted aortic root
24 24 24
29.76 36 25.2
37.2 41.28 32.88
24 24.96 19.68
8.16 10.8 13.44
Fig. 6 Grid pattern im plem ented to verify the absence of opti cal distortions for flow imaging
MAY 2015, Vol. 137 / 0 5 1 0 0 5 -3
Matching the IoR of the experimental model was essential for avoiding optical distortions and achieving excellent image quality for flow visualization. The test fluid used in the current test was as mixture of Nal (sodium iodide) and H20 (water) at the ratio of 1185 g to 697 g by weight. The solution had a viscosity of 0.0032 Pa-s, measured by a calibrated viscometer (Brookfield RVT), and a density of 1650 kg/m3, which is comparable to the viscosity and the density of the blood of 0.003 Pa-s and the density of 1100 kg/m3, respectively. The solution achieved an index of refraction of approximately 1.51, measured by a calibrated refractometer (super scientific, model #300034), resulting in minimal optical distortion as shown in Fig. 6. In addition, a small quantity of Na2S20 3 (sodium thiosulphate) was used to render the solution color less. It is important to note that the effect of the addition of sodium thiosulphate on the physical properties of the fluid was negligible. Neither density viscosity nor IoR was changed by discoloration. The PIV system used in the cardiovascular flow experiments contained a Litron LDY 300 series diode-pumped Nd:YLF dual cavity laser with a maximum energy output of 22.5 mJ at 1 kHz, operating at a wavelength of 527 nm. The laser sheet with a
thickness of 1 mm, measured using the type 1895 linagraph laser bum paper, was generated by a series of cylindrical and spherical lenses. The flow was seeded with silver-coated hollow glass spheres with a mean diameter of 14 /an and specific gravity of 1.3 relative to water. The Stokes number of the droplets was equal to 0.7 x 10 3, which indicated that the particles were sufficiently small to accurately follow the flow [35], The images of the tracers were recorded using a photron complementary metal oxide semi conductor camera with a sensor that consisted of 1024 x 1024 pix els with a pixel size of 17 /an, which was oriented parallel to the data acquisition plane shown in Fig. 3(a). The field of view was 70 mm x 30 mm, which represented a physical region extending i -5D0 upstream and 3D0 downstream of the valve (see Fig. 5). The origin (X = 0 mm and Y = 0 mm) corresponded to the center of the outflow area of the gasket holding the valve. The images were recorded in a double frame/double exposure mode at a rate of 100 image pairs per second and with the time interval between the frames in a pair of 500 /is. The velocity vec tor fields were calculated using commercial PIV software
Plane 1
X, mm
■ i.. . -0 .2 0 0 .2 0
0.60
1.00
1.40
1.80
2.20 -0.20 0.20 0 60
Z, mm
Plane 2
Z, mm
1.00
1 40
1 80 2.20
Z, mm
Fig . 7 S tre a m lin e p a tte rn s a n d c o n to u rs o f v e lo c ity m a g n itu d e (F avg, m /s) a t t / T = 0.13; (a) n o rm a l g e o m e try , lb ) d ila te d a o rtic ro o t, a n d (c ) c o n s tric te d a o rtic ro o t
051005-4 / Vol. 137, MAY 2015
Transactions of the ASME
P eak v a lu e s o f v e lo c ity at t / T - 0 .1 3 and lo c a tio n o f re a tta c h m e n t p o in ts in th re e e x p e rim e n ta l c a ses
T a b le 2
Plane 1 ^max (m/s)
Reattachm ent point (X, Y) (mm)
^max (m/s)
Reattachm ent point (Z, Y) (mm)
2.37 2.51 2.64
(-1 4 ,4 0 ) (-1 6 ,6 4 ). ( -1 4 ,3 4 ) (14,52)
2.45 2.5 2.53
(-1 6 ,3 2 ) (14,44) (-1 6 ,5 6 ) (-1 4 ,2 8 ) (14,38)
Normal Dilated aortic root Constricted aortic root
(a) .
m
Plane 2
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Plane 2 Fig . 8
C o n to u rs o f R S S (d y n /c m 2) at t / T - 0.1 3 ; (a) n o rm a l g eo m etry, (b) d ila te d a o rtic ro o t, an d (c) c o n s tric te d a o rtic root
(LaVision Davis 8.1.3). A multiple-pass cross-correlation algo rithm, with the initial pass of 64 pixel x 64 pixel interrogation windows with a 50% overlap between the adjacent windows in both X- and Y-directions, and the final pass of 16 pixel x 16 pixel interrogation windows with a 50% overlap in both directions was Journal of Biomechanical Engineering
implemented. The resulting vector-to-vector spacing was equal to 8 pixels. Ensemble-Averaging of the Data. In order to provide an accurate description of various properties of the flow, the MAY 2015, Vol. 137 / 051005-5
f,
N
1
1
u,(p Mrms
N
f>/2
=
k=
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N
^ =
^ k=
= v” - v -
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P e te r O s h k a i Mem. ASME Department of Mechanical Engineering, University of Victoria, P.0. Box 1700, STN CSC, Victoria, BC V8W2Y2, Canada
The Influence of the Aortic Root Geometry on Flow Characteristics of a Prosthetic Heart Valve In this paper, performance of aortic heart valve prosthesis in different geometries of the aortic root is investigated experimentally. The objective of this investigation is to estab lish a set of parameters, which are associated with abnormal flow patterns due to the flow through a prosthetic heart valve implanted in the patients that had certain types of valve diseases prior to the valve replacement. Specific valve diseases were classified into two clinical categories and were correlated with the corresponding changes in aortic root geometry while keeping the aortic base diameter fixed. These categories correspond to aortic valve stenosis and aortic valve insufficiency. The control case that corresponds to the aortic root of a patient without valve disease was used as a reference. Experiments were performed at test conditions corresponding to 70 heatslmin, 5 .5 Llmin target car diac output, and a mean aortic pressure of 100 mmHg. By varying the aortic root geome try, while keeping the diameter of the orifice constant, it w a s possible to investigate corresponding changes in the levels of Reynolds shear stress and establish the possibility of platelet activation and, as a result of that, the formation of blood clots. [DOT: 10.1115/1.4029747]
Introduction Prosthetic heart valves such as mechanical (MHV), polymeric (PV), and tissue valves have been used in heart valve replacement over many decades. Successful analysis of the flow through pros thetic heart valves depends on sufficient understanding of the con ditions under which natural valves function. Most disorders of the heart initiate within the left ventricle, as this chamber is subjected to the highest mechanical loads. The flow through the left ventri cle is regulated by the mitral and the aortic valves, which influ ence the inflow and the outflow conditions, respectively [1], The mitral and the aortic valves are the most commonly affected heart valves in a diseased heart, and they are responsible for 34% and 44% of morbidity [2,3], respectively. Much progress has been achieved in the development of artifi cial heart valves [4,5]. Nevertheless, serious problems still exist due to abnormally high-velocity gradients that are present within the jets that emanate from these prosthetic valves. These gradients result in elevated shear stresses that may cause red blood cell damage, platelet activation, and thrombus formation [6,7], which in turn lead to thromboembolism, hemorrhage, and tissue over growth. Flow-induced stresses in blood, acting on a cellular level, have been known to cause thrombus initiation within the compo nents of the mechanical valve prostheses [8], Regions of elevated stresses during the complex motion of the leaflets in some cases lead to structural failure of the MHVs. In vivo and in vitro experi mental studies have yielded valuable information on the relation ship between hemodynamic stresses and the problems associated with the implants [3,9]. When studying such complex phenomenon as blood flow through the heart valve, the choice of the methodology is critical. Two major approaches employed by researchers in the field of
‘Corresponding author. Manuscript received September 3, 2014; final manuscript received February 4, 2015; published online March 5, 2015. Assoc. Editor: Ender A. Finol.
Journal of Biomechanical Engineering
cardiovascular fluid mechanics are experiments and computa tional fluid dynamics (CFD) simulations. The experimental methodology had been used extensively in the investigation of various cardiovascular devices and novel intervention techniques. It is still a method of choice required for quantitative assessment of the hemolytic and thrombogenic poten tial when a novel cardiovascular device is undergoing a regulatory submission for a premarket approval [10]. Among several experi mental methods that allow flow visualization, digital particle image velocimetry (PIV) is particularly useful, as it can deliver global, quantitative flow images with high spatial and temporal resolution. In Refs. [11-14], PIV was used to quantify the velocity field in the vicinity of valve leaflets. The authors of Ref. [15] were among the first to apply PIV methodology to study flow through an artificial heart valve, and they outlined experimental challenges associated with it. The occurrence, severity, and distribution of viscous and Reynolds shear stresses (RSS) were successfully esti mated via PIV. Recently, microcomputed tomography and stereo scopic digital PIV techniques were used to investigate the hemodynamics of the mitral valve [16] and to evaluate a new design of bileaflet mitral valve prosthesis [17]. CFD methodology has been actively employed for the last dec ade to aid in the evaluation of medical devices and to improve safety and quality in the development of products and technolo gies. While still requiring extensive validation, CFD is a very powerful tool that can be successfully used in modeling of com plex, chemico-biological phenomena. The primary advantage of CFD over experimental fluid mechanics is its ability to provide fast analysis of the results of small changes in the design of medi cal devices. The authors in Ref. [18] used unsteady Reynoldsaveraged Navier-Stokes equations simulations in order to assess the flow through a 2D bileaflet MHV under steady inflow condi tions and estimate the influence of the flow field on the platelet activation mechanism. In Refs. [9] and [19], the CFD methodol ogy was applied to investigate complex flows through a bileaflet MHV in an axisymmetric aorta and within the total
Copyright ©2015 by ASME
MAY 2015, Vol. 137 / 051005-1
cavopulmonary connection, respectively. In Ref. [20], a full threedimensional (3D) fluid-structure interaction model of the aortic valve with coaptation and a compliant aorta was used to resolve the kinematics of the aortic valve and the details of leaflet coapta tion at the end of the closing phase of the cardiac cycle. The authors in Refs. [21] and [22] investigated the possible influence of different bifurcation stenting techniques on stent deformation as well as physical stress and drug elution via coupled mechanical-CFD numerical model. The primary consideration for the present study was the devel opment of anatomically correct benchtop experimental model with the ability to investigate the correlation between pathological changes in the geometry of aortic root induced by aortic valve ste nosis and insufficiency and presence of the abnormal flow patterns downstream of the recently implanted valve. The main design considerations for the experimental system included robustness, versatility, and compatibility of the model with the commercial left ventricular heart simulator. In addition to simulating anatomi cally correct geometries, the main physiological features of the healthy valve, such as nonnotensive pulsatile flow and pressure waveforms, were also implemented.
Experimental System and Techniques The experiments described in this paper were performed in the laboratory of ViVitro Labs, Inc., in Victoria. BC, Canada.
configurations while assessing their performance and function under simulated cardiac conditions. The pulse duplicator was powered by a piston-in-cylinder pump driven by a motor connected to the ViViTest data acquisition and control block, which controlled the pump. This setup allowed for various ventricular waveform states and beat rates while generat ing physiological pressures and flow regimes through the heart valve. The present investigation was focused on the hemodynamics of a prototype trileaflet PV, shown in Fig. 3(a). The valve, which had a nominal diameter of 19 mm, was mounted in a model of the aor tic root with severe pathological changes in the aortic root geometry. For each of the aortic root types, the flow fields were collected in two perpendicular data acquisition planes, as shown in Fig. 3(b). The start of the image capture was triggered by a syn chronization pulse from the pump controller. The PIV data were acquired at six representative phases of the cardiac cycle, indi cated by black circles in Fig. 4. These phases correspond to the opening acceleration phase (t/T = 0.05), the peak systole phase (Cl T = 0.13), the deceleration phase (tIT= 0.22), the closing phase (tl T = 0.29), and the leakage phase (t/T = 0.69). Here, t denotes time, and T = 860 ms is the period of the simulated cardiac cycle. Aortic Root Geometry. The aortic root is a part of the ascending aorta, shown in Fig. 5, with a combined length h = 80 mm [1,23,24], In this paper, the aortic root was defined by means of the following parameters [25,26]: D0 is the diameter of the orifice
Pulsatile Flow System. In Figs. 1 and 2, one can see the sche matic of the experimental apparatus that allows testing of the heart valves in the aortic configuration. The ViVitro pulse duplicator is a heart model that incorporates the functional capability to test two heart valve substitutes simultaneously in left or right heart
Fig. 3 (a) Orientation of the valve with respect to the left coro nary artery, the right coronary artery and the noncoronary cusp and the PIV data acquisition planes (dashed lines) and ( b ) sche matic of the prototype trileaflet PV Fig. 1 Schematic of the flow visualization setup (image cour tesy of ViVitro Labs, Inc.)
T im e (m s)
Fig. 2 Schematics of the test chamber and the aortic valve (image courtesy of ViVitro Labs, inc.) 051005-2
/ Vol. 137, MAY 2015
Fig. 4 Variation of flow rate as a function of time during a typi cal cardiac cycle. Black circles correspond to the phases of the cardiac cycle, at which PIV data were obtained.
Transactions of the ASME
Fig. 5 M ajor parts of the aorta and principal dim ensions of the aortic root sinuses
(also referred as aortic annulus diameter), DA is aortic diameter distal to sinus or diameter of the sinutubular junction, DB is the maximum projected sinus diameter, which approximate the diam eter of anatomic ventriculoarterial junction, LA is the length of the sinuses, and LB is the distance between D0 and DB. It should be noted that Da , the diameter of the orifice, was kept constant in order to interface properly with the ViVitro pulse duplicator. The primary objective of this study was to investigate the path ological changes in the dimensions of the aortic root due to aortic valve disease, such as valve stenosis and valve insufficiency, and to determine the influence of those changes on the appearance of abnormal flow patterns in the flow through the aortic valve which replaced the original, diseased valve. These pathological changes in extreme scenarios are often associated with patients morbidity due to aortic dissection, aortic incompetence caused by a dilated, aneurysmal aortic root, and severe aortic valve stenosis [27-30], In addition, deformation of the aortic root after valve replacement or structural dysfunction of the recently replaced bioprosthetic heart valve due to premature calcification associated with pure stenosis due to cusps stiffening is not uncommon [31,32]. Proper replication of the complex anatomy of the aortic valve and root is essential in many ways. First, the accurate measure ment of the aortic root geometry in clinical settings is required for evaluation of the patients with aortic valve or aortic root disease and determining whether the proposed device can be safely and successfully implanted, which is primarily related to the newly emerged transcatheter aortic valve implantation (TAVI) [25,27,33]. Second, the accurate representation of a complex aor tic root anatomy is essential in order to reproduce the internal physiological flow field correctly for any in vitro study. During just one cardiac cycle that includes the systole and the diastole phases, several complex flow events occur within the aortic root. The opening and the closing motions of the valve are supported through vortex formation within the aortic sinuses. In addition, when the valve is closed, previously formed sinus vortices are responsible for washout of the sinus cavities, so that the throm botic deposition does not occur. Among many techniques that are capable of a nonintrusive imaging of elements of the cardiovascular system, angiography is probably the most well known. The projectional radiography or angiography is a technique that employs a special dye in order to visualize the blood flow in arteries as well as the dimensions of the arteries. Nowadays, there are several highly sophisticated methodologies that are capable of accurate assessment of the major dimensions of the aortic root [25,27,34]. One of the most commonly used techniques for such assessment is transthoracic echocardiography (TTE). TTE can be applied to determine the
Journal of Biomechanical Engineering
severity of aortic valve/root disease and whether the patient is suitable for a TAVI procedure [27]. In addition to TTE, a contrast-enhanced multidetector computed tomography (MDCT), due to its accuracy, can be applied to determine the difference in aortic root geometry between patients with tricuspid and bicuspid aortic valve. Among the major drawbacks of MDCT is the exces sive radiation exposure to the patient and the risk of hypersensitiv ity to iodine contrast. As an alternative to MDCT, a 3D transesophageal echocardiography (TEE) has emerged and been demonstrated reliably to evaluate aortic root geometry [34]. The importance, of studying aortic root geometry associated with a particular clinical case in conjunction with a valve pros thetic in a single study, was stressed in Ref. [26], The authors showed a concise quantification of the aortic root geometry changes due to valvular diseases. Geometric parameters and alge braic quotients were obtained from 604 independent angiographic films collected from patients with various types and degrees of valvular diseases [26]. These calculated quantities were used in the present work in such way that the resulting aortic root geome try was compliant with the existing ViVitro pulse duplicator sys tem. By setting the orifice diameter to D0 = 24 mm and the wall thickness to A = 1 .5 mm, the dimensions listed in Table 1 were derived from relationships originally presented in Ref. [26]. In this paper, the geometries of the aortic root that corresponds to the aortic valve stenosis and aortic valve insufficiency [26] are referred as dilated aortic root and constricted aortic root, respectively. Based on the dimensions presented in Table 1, three types of geometries that represented aortic root and ascending aorta system were built through the process of stereolithography (SLA). DSM Somos® Watershed© XC 11122 (Watershed), an optically clear resin with acrylonitrile butadiene styrene-like properties, good tem perature resistance, water resistance, and high durability was used as a base material. This material is also a Class VI Approved Medi cal Grade material making it suitable for any medical applications. Watershed material has an index of refraction (IoR) of 1.51. Table 1 Principal dim ensions of the aortic root associated with clinical cases of heart disease
D0 (mm) Da (mm) DB (mm) LA (mm) LB (mm) Normal Dilated aortic root Constricted aortic root
24 24 24
29.76 36 25.2
37.2 41.28 32.88
24 24.96 19.68
8.16 10.8 13.44
Fig. 6 Grid pattern im plem ented to verify the absence of opti cal distortions for flow imaging
MAY 2015, Vol. 137 / 0 5 1 0 0 5 -3
Matching the IoR of the experimental model was essential for avoiding optical distortions and achieving excellent image quality for flow visualization. The test fluid used in the current test was as mixture of Nal (sodium iodide) and H20 (water) at the ratio of 1185 g to 697 g by weight. The solution had a viscosity of 0.0032 Pa-s, measured by a calibrated viscometer (Brookfield RVT), and a density of 1650 kg/m3, which is comparable to the viscosity and the density of the blood of 0.003 Pa-s and the density of 1100 kg/m3, respectively. The solution achieved an index of refraction of approximately 1.51, measured by a calibrated refractometer (super scientific, model #300034), resulting in minimal optical distortion as shown in Fig. 6. In addition, a small quantity of Na2S20 3 (sodium thiosulphate) was used to render the solution color less. It is important to note that the effect of the addition of sodium thiosulphate on the physical properties of the fluid was negligible. Neither density viscosity nor IoR was changed by discoloration. The PIV system used in the cardiovascular flow experiments contained a Litron LDY 300 series diode-pumped Nd:YLF dual cavity laser with a maximum energy output of 22.5 mJ at 1 kHz, operating at a wavelength of 527 nm. The laser sheet with a
thickness of 1 mm, measured using the type 1895 linagraph laser bum paper, was generated by a series of cylindrical and spherical lenses. The flow was seeded with silver-coated hollow glass spheres with a mean diameter of 14 /an and specific gravity of 1.3 relative to water. The Stokes number of the droplets was equal to 0.7 x 10 3, which indicated that the particles were sufficiently small to accurately follow the flow [35], The images of the tracers were recorded using a photron complementary metal oxide semi conductor camera with a sensor that consisted of 1024 x 1024 pix els with a pixel size of 17 /an, which was oriented parallel to the data acquisition plane shown in Fig. 3(a). The field of view was 70 mm x 30 mm, which represented a physical region extending i -5D0 upstream and 3D0 downstream of the valve (see Fig. 5). The origin (X = 0 mm and Y = 0 mm) corresponded to the center of the outflow area of the gasket holding the valve. The images were recorded in a double frame/double exposure mode at a rate of 100 image pairs per second and with the time interval between the frames in a pair of 500 /is. The velocity vec tor fields were calculated using commercial PIV software
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Fig . 7 S tre a m lin e p a tte rn s a n d c o n to u rs o f v e lo c ity m a g n itu d e (F avg, m /s) a t t / T = 0.13; (a) n o rm a l g e o m e try , lb ) d ila te d a o rtic ro o t, a n d (c ) c o n s tric te d a o rtic ro o t
051005-4 / Vol. 137, MAY 2015
Transactions of the ASME
P eak v a lu e s o f v e lo c ity at t / T - 0 .1 3 and lo c a tio n o f re a tta c h m e n t p o in ts in th re e e x p e rim e n ta l c a ses
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C o n to u rs o f R S S (d y n /c m 2) at t / T - 0.1 3 ; (a) n o rm a l g eo m etry, (b) d ila te d a o rtic ro o t, an d (c) c o n s tric te d a o rtic root
(LaVision Davis 8.1.3). A multiple-pass cross-correlation algo rithm, with the initial pass of 64 pixel x 64 pixel interrogation windows with a 50% overlap between the adjacent windows in both X- and Y-directions, and the final pass of 16 pixel x 16 pixel interrogation windows with a 50% overlap in both directions was Journal of Biomechanical Engineering
implemented. The resulting vector-to-vector spacing was equal to 8 pixels. Ensemble-Averaging of the Data. In order to provide an accurate description of various properties of the flow, the MAY 2015, Vol. 137 / 051005-5
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