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Accepted Manuscript
Finding the Ideal Biomaterial for Aortic Valve Repair with ex-vivo Porcine Left HeartSimulator and Finite Element Modeling
Hadi Daood Toeg, MD, MSc Ovais Abessi, MEng Talal Al-Atassi, MD, MPH Laurentde Kerchove, MD Gebrine El-Khoury, MD Michel Labrosse, PhD Munir Boodhwani,MD, MMSc
PII: S0022-5223(14)00525-X
DOI: 10.1016/j.jtcvs.2014.05.004
Reference: YMTC 8590
To appear in: The Journal of Thoracic and Cardiovascular Surgery
Received Date: 21 February 2014
Revised Date: 13 April 2014
Accepted Date: 2 May 2014
Please cite this article as: Toeg HD, Abessi O, Al-Atassi T, de Kerchove L, El-Khoury G, Labrosse M,Boodhwani M, Finding the Ideal Biomaterial for Aortic Valve Repair with ex-vivo Porcine Left HeartSimulator and Finite Element Modeling, The Journal of Thoracic and Cardiovascular Surgery (2014), doi:10.1016/j.jtcvs.2014.05.004.
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Finding the Ideal Biomaterial for Aortic Valve Repair with ex-vivo
Porcine Left Heart Simulator and Finite Element Modeling
Hadi Daood Toeg1 MD, MSc
Ovais Abessi2 MEng
Talal Al-Atassi1 MD, MPH
Laurent de Kerchove3 MD
Gebrine El-Khoury3 MD
Michel Labrosse2 PhD
Munir Boodhwani1 MD, MMSc
1. Division of Cardiac Surgery, University of Ottawa Heart Institute, Ottawa, Ontario, Canada
2. Department of Mechanical Engineering, University of Ottawa, Ottawa, Ontario, Canada
3. Department of Cardiovascular and Thoracic Surgery Cliniques Universitaires Saint-Luc,
Brussels, Belgium
Short Title (46/50c): Finding the ideal biomaterial for aortic valve repair
Journal of Thoracic and Cardiovascular Surgery; Original Manuscript
Corresponding Author: Dr. Munir Boodhwani ([email protected])
40 Ruskin Street, Ottawa, Ontario, Canada
Phone: 613-761-4720 Fax: 613-761-4713
Word count: 3459/ 3500 max (excludes references, legends, abstract)
Figures & tables: 5/5 max (2 additional in online supplement)
References: 30/35max
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Abstract (250/250):
Objectives: Aortic valve repair (AVr) has become an attractive alternative for the
correction of aortic insufficiency as compared to aortic valve replacement; however, little
clinical evidence exists in determining which biomaterial at AVr is optimal. Cusp
replacement in AVr is associated with increased long-term repair failure. We measured
hemodynamic and biomaterial properties with the use of an ex-vivo porcine AVr model
with clinically relevant biomaterials and generated a finite element model to ascertain
which material(s) would be best suited for valve repair.
Methods: Porcine aortic roots with intact aortic valves were placed in a left heart
simulator mounted with a high-speed camera for baseline valve assessment. The non-
coronary cusp (NCC) was excised and replaced with autologous porcine pericardium
(APP), glutaraldehyde-fixed bovine pericardial patch (BPP; Synovis™), extracelluar
matrix scaffold (CorMatrix™), or collagen-impregnated Dacron (HEMASHIELD™).
Hemodynamic parameters were measured over a range of cardiac outputs (2.5–
6.5L/min) post-repair. Biomaterial properties, along with St. Jude Medical™ Pericardial
Patch (SJM), were determined using pressurization experiments; consequently, finite
element models of the aortic valve and root complex were constructed to determine
hemodynamic characteristics and leaflet stresses.
Results: Post-repair geometric orifice areas (GOAs) were significantly reduced in the
HEMASHIELD™ (P<0.05), and CorMatrix™ (P=0.0001) groups. Left ventricular work
(LVW) increased with increasing cardiac output (P=0.001) in unrepaired valves as
expected, and was similar between all biomaterial groups. Finite element modeling of
biomaterials displayed differences in percent changes in total Von Mises stress for both
replaced (NCC) and non-replaced left and right cusps with SJM (+4%, +24%) and APP
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(+5,+26%) having lower percent changes than BPP (+12%, +27%), HEMASHIELD™
(+30%, +9%), and CorMatrix™ (+13%, +32%).
Conclusions: This study demonstrates that post-repair LVW did not increase despite a
drop in GOAs in the HEMASHIELD™ and CorMatrix™ groups. APP and SJM have the
closest profile to normal aortic valves; therefore, use of either biomaterial may be best
suited. Finally, increased stresses found in BPP, HEMASHIELD™, and CorMatrix™
groups may after prolonged tensile exposure be associated with late repair failure.
Keywords:
Aortic surgery, modeling, valve dynamics
Ultramini Abstract (46/50w):
With the use of a left heart simulator model and finite element modeling,
autologous porcine pericardium and the St. Jude Medical bovine pericardial patch
demonstrated the closest hemodynamic profile to normal aortic valves; therefore, use of
either biomaterial may be best suited for aortic valve repair.
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List of Abbreviations:
AI: aortic insufficiency
APP: autologous porcine pericardium
AV: aortic valve
AVr: aortic valve repair
BPP: bovine pericardial patch
CO: cardiac output
CorM: CorMatrix
FEM: finite element model
GOA: geometric orifice area
HEM: Hemashield
LHS: left heart simulator
LVW: left ventricular work
NCC: non-coronary cusp
SJM: St. Jude Medical bovine pericardial patch
VC: valve closing velocity
VO: opening velocity
VSC: valve slow closing velocity
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Introduction:
Aortic valve repair (AVr) has evolved into an attractive strategy for the correction
of aortic insufficiency (AI) compared to valve replacement1. Successful repair can be
achieved if the surgeon addresses the specific anatomical defects with high surgical
flexibility, knowledge, and adequate surgical tools in his armamentarium2. Attempts to
replace cusp tissue with biological material have been made since the 1960s with
materials such as dura mater, facia lata, and bovine pericardium3. Several groups have
attempted to mimic autologous aortic valve tissue with biomechanical engineering
strategies like tissue bioreactors4, while others studied commercially available
biomaterials in various in vitro and in vivo settings5-7.
Glutaraldehyde-treated bovine pericardial patch and either treated or untreated
autologous human pericardial tissue are common biomaterials used in both cardiac and
vascular surgery1, 8, 9. Patch repair may be used in a variety of settings. These include
repair of cusp perforation due to healed endocarditis or for cusp restoration following
resection of a calcified or restrictive raphe of a bicuspid aortic valve. A patch may also
be used to improve cusp coaptation in pediatric AVr or to reconstruct a commissure
when bicuspidizing a unicuspid or tricuspidizing a bicuspid aortic valve 8.
An increased risk of long-term recurrent AI has been associated with the use a
patch in a number of studies8, 10 utilizing a variety of patch materials and may be due to
the limitations of existing biomaterials available for cusp reconstruction. Thus, while the
perfect biomaterial remains elusive, we sought to ascertain which material(s) would be
best suited for cusp restoration in AVr. By utilizing an AV-cusp restoration ex vivo
porcine left heart simulator model, we measured hemodynamics, biomaterial properties,
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and developed a finite element model (FEM) comparing native porcine aortic valves to
repaired valves using clinically relevant biomaterials.
Methods:
Ex Vivo Porcine Aortic Valve Repair Model:
Fresh aortic roots (N=25), along with 6 cm of ascending aorta, were dissected
out from healthy porcine (donated in kind by a local commercial supplier) and coronary
arteries were ligated. All aortic roots were connected to the ViVitro Left Heart Simulator
(LHS; ViVitro Systems Inc., Victoria, Canada) for baseline valve measurements acting
as paired controls. After baseline recordings, a transverse aortotomy was created 1 cm
above the sinotubular junction and the non-coronary cusp (NCC) was resected, leaving
a 3 mm cuff of tissue attached to the aortic wall. An outline of the excised cusp was used
to create a 2 mm oversized cusp tissue from various biomaterials. The biomaterial was
then connected, at both lateral edges, with two 6-0 polypropylene sutures (Ethicon,
Johnson & Johnson Inc., Montreal, Canada). Suturing of the patch was performed using
a continuous running suture. The aortotomy was closed in double layers with 5-0
polypropylene (Ethicon).
Biomaterials:
Four clinically relevant materials were used in this study. The first material was
derived from fresh porcine pericardial tissue taken from the same porcine heart to which
the valve repair was performed. This was kept in 0.9% normal saline. This biomaterial
will be referred to as autologous porcine pericardium (APP). Next, a collagen
impregnated double woven Dacron HEMASHIELD patch™ was used (HEM; MAQUET,
Rastatt, Germany). Next material utilized was CorMatrix™ (CorM; CorMatrix™, Roswell,
USA): a biomaterial derived from porcine small intestinal submucosa and used for repair
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of certain cardiac defects11. Finally, with over 30 years of cardiovascular implantation
experience, the PERI-GUARD bovine pericardial patch was used (BPP; Synovis
Surgical Innovations, Deerfield, IL, USA). All biomaterials above were used in the LHS
model. In addition to the biomaterials listed above, we considered another bovine
pericardial patch donated by St. Jude Medical™ (SJM, St. Jude Medical Inc., St. Paul,
MN, USA) to perform our finite element modeling experiments (see ‘Finite Element
Model’ & ‘Simulation of Leaflet Repair’ sections). The SJM™ Pericardial Patch with
EnCap™ AC Technology (SJM) is a thinner, flexible bovine pericardial patch with a
different anti-calcification treatment compared to BPP15(supplementary Table 2).
Left Heart Simulator Measurements:
The Vivitro LHS was used as a pulse duplicator with capabilities of measuring
flow, pressure, and generating a wide range of cardiac outputs. This LHS consists of a
piston pump system, a visco-elastic impedance adapter for afterload generation, an
adjustable left ventricular membrane, flow and pressure monitoring systems, a waveform
generator, and a data acquisition system. In addition, an adjustable jig was used to
attach the repaired aortic roots to the system. A high-speed camera (Phantom V4.2,
Vision Research, Wayne, USA) was connected to an endoscope and light source to
record images of valvular leaflet motion from the aortic side. Further details are
described previously from our group12.
For valve hemodynamic measurements, the unrepaired or repaired aortic root
was connected at room temperature to the LHS. The system, containing 0.9% normal
saline (density of 0.9 g/ml and a viscosity of 1 mPa-s) as the fluid medium, was
pressurized from 0 mmHg to a systolic pressure of 120 mmHg and diastolic pressure of
80 mmHg. The heart rate was set at 70 for all experiments, while cardiac output (CO)
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was varied from 2.5 to 6.5 L/min. Parameters measured include: valve opening velocity
(VO), valve slow closing velocity (VSC), valve closing velocity (VC), maximum geometric
orifice area (GOA), and left ventricular work (LVW). The AI fraction was insignificant in
all treated aortic roots.
Finite Element Model:
Following the procedure outlined in our previous study12, a FEM of a typical
unpressurized porcine aortic valve with root was built with dimensions listed in the online
supplementary tables 1 and 2. The structured hexahedral FEM mesh consisted of
approximately 11,800 nodes and 8,600 elements. The valve model was studied over one
cardiac cycle by application of set pressure pulses12. Since the analysis started from the
unpressurized geometry, the pressure was ramped from 0 mmHg up to 80 mmHg before
the physiological cardiac cycle started in early systole. While the normal duration of
diastole is approximately 0.60 s, it was shortened to 0.10 s in the simulation to save
computational time. In addition, the simulation time was 1/10 of the real time, as
analyses with real or scaled time yielded results within 2% of each other, due to
comparatively small inertial loads. All numerical analyses were carried out with
commercial FEM software LS-Dyna 971 (LSTC, Livermore, CA, USA). The mass density
of the aortic and leaflet tissues was set at 1,000 kg/m3. The time step was automatically
set and updated by LS-Dyna to achieve numerical stability of the solution. The
longitudinal stretch ratio was set to the normal physiological value of 1.2013. Tissues of
the aorta and of the aortic leaflets were both modeled as hyperelastic, transversely
isotropic materials using a Fung-like model with strain energy function
� = ��� �exp���
� + ������ + ��� + ���� + ���� � + ���� �� + � �� +�� � + �� � �� − 1� +�� �� − 1�,
where c1, …, c4 are material constants, E…are deformations (Green strain components
modified to only include the effects of volumetric work), and subscripts θ, z, r refer to the
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circumferential, longitudinal and radial directions, respectively. P is a Lagrange multiplier
numerically enforcing the material near-incompressibility whereby J, the determinant of
the deformation gradient tensor, is almost equal to 1. Although this material model was
initially developed to represent myocardial tissue, it has been shown to give accurate
representations of human aortic and leaflet tissues14. The material constants for the
original porcine model are listed in supplementary table 2. Unknown constant c4 in the
material model was assumed to equal c3 – see14 for justification.
Simulation of Leaflet Repair:
With all other parameters unchanged, the thickness and material properties of
the non-coronary leaflet (NCC) in the original porcine valve model were modified to
represent those of the biomaterials considered in this study. Biomaterial properties were
measured with static pressurization experiments. Each sample was sutured into a
cylinder, cannulated at one end, capped at the other, and placed in normal saline at
room temperature. The sample was held horizontally by the cannulated end which was
connected to a saline solution head whose pressure was measured by a digital
manometer (33500-082, VWR International, Mississauga, ON, Canada). The sample
was preconditioned by three one-minute pressurizations from 0 mmHg to 160 mmHg to
0 mmHg. The elongation of the sample was monitored using a digital camera
(Powershot IS2, Canon, Tokyo, Japan) and two markers affixed to it. For the
measurement run, pressure was slowly stepped up from 0 mmHg to 160 mmHg in
increments of 20 mmHg. Cross-sectional views of the sample being loaded were
acquired using an ultrasound external probe (Terason SMARTProbe, 5-8 MHz 4-cm
linear array “T” probe, Teratech Corp., Burlington, MA, USA) driven by the PC-based
TelaVet 1000 system (Classic Medical, Tequesta, FL, USA). Corresponding images
synchronized with the pressure steps were stored. Processing of similar data has been
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previously detailed14; herein, it yielded the properties listed in supplementary table 2.
Figure 1 illustrates both circumferential and radial direction stress-strain relationships for
various biomaterials (colour coded) under equibiaxial planar testing driven in tension. In
all the FEMs, the maximum values of mechanical Von Mises stress in the leaflets were
determined, along with their location and timing. Von Mises stress summarizes in one
number the stresses present in different directions of the material at one specific location.
In addition, to evaluate the VO and VS, the GOA of the valve was measured as a
function of time from the calculated leaflet motions throughout one cardiac cycle. This
was done by using the post-processor of LS-Dyna to take top view snapshots of the
valve at regular time intervals during the cardiac cycle and by measuring the projected
area left open by the leaflets.
Statistical Analysis:
Statistical analysis was performed with STATA v12 statistical program
(StataCorp LP, College Station, Texas). Nonparametric data are expressed as the
median value with interquartile range. Wilcoxon-signed rank tests were used to
determine significance between paired sample sets while Kruskal-Wallis rank test was
used for more than two sample sets. Significance was set at 0.05.
Results:
A total of 25 fresh porcine aortic roots underwent baseline measurements. Aortic
valves subsequently underwent randomized allocation to NCC excision and cusp
replacement with APP (N=7), HEM (N=7), CorM (N=5) and BPP (N=6). Each unrepaired
and repaired aortic valve underwent 3 sample runs at set cardiac outputs of 2.5, 5.0, and
6.5 L/min.
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Hemodynamic Results from LHS:
For appropriate comparisons, each treatment sample was compared to its
original valve measurements (paired experiments). Compared to baseline valve
hemodynamics, the APP biomaterial demonstrated significantly lower VO (P=0.02), and
VC values (P=0.03) (Table 1). Although not statistically significant, post repair GOA was
reduced from 5.74 cm2 (5.29 - 6.87 cm2) to 5.25 cm2 (4.16 - 5.99 cm2) (P=0.08). Despite
this drop in GOA, the LVW post-repair did not change (P=0.6). After replacing NCCs with
HEM biomaterial, there was a significant reduction in the GOA (5.91 cm2 [5.58 - 6.43
cm2] to 5.06 cm2 [4.24 - 5.73 cm2]; P=0.05) and VC (63 cm/s [59 - 71 cm/s] to 48 cm/s
[35 - 61 cm/s]; P=0.01) compared to baseline (Table 1). All other parameters including
VO, VSC, and LVW did not significantly change (P>0.3). Post-repair valve
hemodynamics with CorM demonstrated significant reductions in VO (P = 0.0001), VSC
(P=0.01), VC (P=0.0001), and GOA (P=0.0001; Table 1). Despite these findings, the
LVW was similar between baseline and repaired valves (1033 mJ [694 - 1221 mJ] to
1086 mJ [727 - 1241 mJ]; P=0.4). Finally, after comparing BPP-repaired valves to
baseline valve hemodynamics, no significant differences were observed in VO (P = 0.4),
VSC (P=0.08), VC (P=0.8), GOA (P=0.5), and LVW (P=0.4; Table 1).
As expected, when native non-repaired porcine valve hemodynamics at CO
settings of 2.5, 5.0, and 6.5 L/min were compared, LVW and VO increased with
increasing CO (P=0.001, P=0.002, respectively; Table 2). All other parameters, including
GOA were not statistically different (P>0.05; Table 2).
Finite Element Model:
All FEMs opened and closed properly. The hemodynamics (VO, VC) of the
repaired valves were not distinguishable from those of the original valve (Table 3). The
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repaired valves opened and closed more slowly than the original valve, by at most 9%.
The GOAs of the repaired valves were between 6 – 9% smaller than that of the original
unrepaired valve. In the repaired valves, all biomaterials produced different maximum
Von Mises stresses in all leaflets (NCC and left/right non-replaced coronary cusps)
(Table 3). The replaced leaflets (NCC) with the smallest stresses were made of SJM and
APP. In the NCC, stresses in SJM and APP were 4% and 5% higher than the stresses in
the native leaflets, respectively. Conversely, the replaced leaflets that experienced the
highest stresses consisted of HEM (+30%), CorM (+13%), and BPP (+12%).
Interestingly, due to the interaction between AV cusps during coaptation, the stresses in
the non-replaced leaflets were also affected by the material utilized. The smallest
increases in stress experienced by the native leaflets (left/right non-replaced cusps)
were 9% when HEM was used for the replaced leaflet, 24% with SJM, and 26% with
APP. The highest increases in stress were 32% with CorM, and 27% with BPP. Finally,
after combining the Von Mises stresses of both replaced (NCC) and non-replaced cusps
(left/right non-replaced cusps), SJM biomaterial had the lowest increase in stress at
+28% (3,968 kPa) followed next by APP with +31% (4,047 kPa) compared to the native
unrepaired model. The highest increased combined Von Mises stress was found in the
CorM group (+45%; 4,282 kPa) followed next by the HEM (+39%; 4,158 kPa), and BPP
groups (+ 39%; 4,158 kPa). An illustration of the Von Mises stresses experienced on the
AV cusps is depicted in Figure 2.
Discussion:
Given the limited durability of AVr in the setting of patch replacement for cusp
tissue, our objective was to determine which biomaterial(s) would be best suited for AV
cusp repair. Utilizing an ex vivo porcine aortic valve LHS model to assess valvular
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hemodynamics and finite element modeling to assess leaflet stresses, we compared a
variety of clinically relevant biomaterials. We found that post-repair LVW did not increase
despite a drop in GOAs in the HEMASHIELD™ and CorMatrix™ groups. Our ex vivo
LHS model suggested that BPP would be best suited for AVr. However, finite element
modeling demonstrated increased cusp stresses in BPP, HEMASHIELD™, and
CorMatrix™ groups, which after a significant amount of exposed force may be
associated with late repair failure. Autologous pericardium and the thinner SJM
pericardial patch had the closest profile to normal aortic valves and would be best suited
for AV cusp repair.
Although clinical experience has suggested increase rates of recurrent AI in
repaired aortic valves when patch material is used for reconstruction8, 16, translating this
finding to a clinically relevant pre-clinical model poses significant challenges. Given the
nature of AV cusp repair, a large animal model would be required as small animal
models would be associated with significant technical limitations. The absence of large
animal models of aortic valve disease, the challenges in performing complex aortic valve
interventions in these animals, and the need to survive these animals for months (if not
longer) to observe clinically relevant changes are some of these challenges. As such,
models that provide detailed information on valve dynamics complemented with data on
valve stresses are the best surrogate indicators available to ascertain long-term
performance of cusp repair materials; moreover, similar approaches have been used in
the development of prosthetic valves17. We therefore, utilized such a combination in this
study.
Clinical applications of autologous human pericardial tissue have been used for
AVr, aortic root enlargement, and LV aneurysm repair1, 9, 10, 18. Since autologous porcine
pericardium (APP) has characteristics similar to human pericardial tissue, APP was used
in this study as a surrogate to human pericardium19. The APP used was derived from
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fresh porcine pericardium and exhibited poor handling qualities due to its fragile
structure. Post-repair VO (P=0.03) and VC (P=0.02) were significantly reduced as
compared to unrepaired valves; however, the GOA and LVW were unchanged. Thus,
APP would be a reasonable choice of biomaterial to use since it demonstrates similar
native valvular hemodynamics. Next, when the HEMASHIELD patch was used to
replace the NCC, VC (P=0.01) and GOA (P<0.05) were significantly reduced while other
parameters including LVW, VSC, and VO were similar. This suggests that despite the
lack of increased LVW, the restricted noncompliant nature of the HEMASHIELD material
lead to slower VC and smaller GOAs. While HEMASHIELD has been used for aortic root
enlargement, LV aneurysm repair, and is also known for its mechanical strength20 and
long-term durability21, it has not been used for AVr in the clinical setting22 and should
likely remain this way based on these results.
With the use of CorM for AVr, all valvular hemodynamic parameters including VO
(P=0.0001), VSC (P=0.01), VC (P=0.0001), and GOA (P=0.0001) were significantly
reduced when compared to baseline. There was a slight increase in LVW post CorM
repair but it did not reach significance (P=0.4). Based on all these early hemodynamic
measurements, this biomaterial may not be the most ideal product for AVr. CorMatrix is
an approved sheet for the use in cardiac surgical septal repair23, and pericardial
closure24, but little information is provided in the literature reporting on clinical application
for AVr11. In addition, our ex-vivo and FEM evaluated CorMatrix prior to autologous cell
seeding, which may alter its biomaterial properties and may not represent long-term
repair characteristics. After repair with BPP valvular hemodynamics did not change
across all parameters (VO, VSC, VC, GOA, and LVW). Along with easy handling
properties, this suggests that BPP is a reasonable choice for biomaterial selection in AVr.
Some studies demonstrate good short-term outcomes following cusp restoration with
bovine pericardial patch25; however, long term recurrent AI was demonstrated in a
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predominately porcine pericardial patch technique used for cusp restoration8, 16. Also,
fresh and photoFixed autologous pericardium trended toward better durability than
glutaraldehyde fixed bovine pericardium but predictors of re-operation did not include the
biomaterial used in their study (a pediatric population)26. Alternatively, one group
demonstrated that the use of a pericardial patch at AVr was a statistically significant
predictor of re-operation further complicating the choice of material to use27.
The valvular hemodynamic measurements based on our ex vivo LHS model
suggests that APP and BPP would be the best suited biomaterial candidates for AVr;
however, no information was available regarding the stress and strain experienced by
the AV cusps. Our FEM demonstrated that while the VO, VC, and GOA values did not
differ by more than 9% (range: 0 – 9%) between all biomaterials and un-repaired valves,
the Von Mises stresses varied significantly between biomaterials in both the NCC
(replaced cusp) and the left/right non-replaced cusps. Von Mises stress represents a
combination of all stresses acting on a particular site28. This variable has been measured
in previous studies in determining specific locations (i.e. the aortic commissures) with
highest dynamic and static stresses12, 14, 29. By understanding the values and patterns of
valvular stresses, surgeons could utilize better, more stress-compatible biomaterials in
the reparative process thereby reducing the incidence of AVr related failure8, 10, 16, 30. The
SJM bovine pericardial patch resulted in the least increase in NCC (+4%; 1,809 kPa)
and left/right non-replaced cusp stresses (+24%; 2,159 kPa). Following closely, the APP
demonstrated the next least increase in NCC (5%; 1,841 kPa) and left/right non-replaced
cusp stresses (+26%; 2,206 kPa). Other groups including HEMASHIELD, CorM, and
BPP demonstrated combined increased stresses ranging from 39 – 45% compared to
baseline. Thus, based on our FEM, the SJM and APP represent the most ideal
biomaterials for AVr. Although the ex vivo data suggests that BPP would be best suited
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for AVr, finite element modeling determined that total cusp stresses was second highest
in this group and that it may, in fact, not be the ideal biomaterial.
In this study, we utilized a model that studies detailed early hemodynamic
parameters following valve repair. The impact of remodeling of the biomaterial over time
on cusp function was therefore, not evaluated. Future long-term experiments are
warranted to establish direct correlation between early hemodynamic profiles on various
biomaterials and their failure rate. Furthermore, other factors including long-term
biocompatibility of leaflet materials and inherent predisposition to calcification or
degradation were not measured, thereby limiting our final conclusions. Other limitations
include the lack of an in vivo model to evaluate post-repair hemodynamics and utilization
of a mathematical FEM to derive stress patterns. While an in vivo model could provide
more realistic settings, it would be difficult to perform the precise measurements of valve
function utilized in this study. Furthermore, the FEM utilized has been used extensively
in the cardiac surgical literature in providing information not readily available from in vivo
studies14, 29.
Altogether, this study demonstrates that post-repair LVW did not increase despite
a drop in GOAs in the HEMASHIELD™ and CorMatrix™ groups. Although the ex vivo
LHS model suggested that BPP would be best suited for AVr, finite element modeling
demonstrated increased stresses in BPP, HEMASHIELD™, and CorMatrix™ groups
which may after prolonged exposure to additional forces, be associated with late repair
failure. Finally, APP and SJM had the closest profile to normal aortic valves in both
models of early valvular hemodynamics; therefore, use of either biomaterial may be best
suited.
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Acknowledgements: We would like to thank CorMatrix and St. Jude Medical for
donating their respective materials. M.L. gratefully acknowledges the Natural Sciences
and Engineering Research Council of Canada for Discovery Grant 312065-2012.
Source of Funding: This work was partially supported by the Natural Sciences and
Engineering Research Council of Canada for Discovery Grant 312065-2012 (M.L.).
Disclosures: None
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Tables and Legends:
Table 1: Comparison of unrepaired (baseline controls) and repaired valve
hemodynamics for biomaterials tested with combined cardiac outputs of 2.5, 5.0 and 6.5
L/min. Median with interquartile ranges are presented.
APP (N=21): autologous porcine pericardium, HEM (N=21): HEMASHIELD, CorM
(N=15): Cormatrix, BPP (N=18): bovine pericardial patch, VO: valve opening velocity,
VSC: valve slow closing velocity, VC: valve closing velocity, GOA: maximum geometric
orifice area, LVW: left ventricular work.
Table 2: Valve hemodynamics in un-repaired porcine valves at different cardiac
outputs (CO).
VO: valve opening velocity, VSC: valve slow closing velocity, VC: valve closing
velocity, GOA: maximum geometric orifice area, and LVW: left ventricular work.
Table 3: Summary of FEM simulation results.
NCC: maximum Von Mises stress in the replaced non-coronary cusp; L/R cusp:
maximum Von Mises stress in the left and right non-replaced coronary cusps, APP:
autologous porcine pericardium, HEM: HEMASHIELD, CorM: Cormatrix, BPP: bovine
pericardial patch, VO: valve opening velocity, VC: valve closing velocity, GOA: maximum
geometric orifice area
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Figure Legends:
Figure 1: Representative circumferential and radial stress-strain properties of the
original leaflets and corresponding biomaterials (see legend with colour scheme) derived
from equibiaxial planar testing (Cauchy stress, kPa; Green strain, %). Original: native
unrepaired valve, Rad: radial direction, Circ: circumferential direction, APP: autologous
porcine pericardium, HEM: HEMASHIELD, CorM: Cormatrix, BPP: bovine pericardial
patch, SJM: St. Jude Medical patch.
Figure 2: Illustration of the Von Mises stresses felt on the replaced cusps (NCC) with
finite element modeling demonstrating A) native unrepaired valve, B) APP, and C)
HEMASHIELD. Blue areas represent minimal to no stress (0 kPa), while light blue, green,
and yellow colours represent increasing levels of stress (colour legend; max 7.774 kPa).
Increased maximal stresses are seen at the suture lines in replaced NCCs for B) APP,
and C) HEMASHIELD compared to A) native unrepaired valve. NCC: non-coronary cusp,
APP: autologous porcine pericardium, HEM: HEMASHIELD.
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Supplementary Material (2 Tables):
Supplementary Table 1: Unpressurized porcine aortic valve dimensions (mm).
LVOT: diameter of left-ventricular outflow tract, STJ: diameter of sinotubular junction, H:
valve height, RSM: maximum radius of aortic sinuses, LH: leaflet height, FE: leaflet free
margin, Hs: sinus height, Hc: height of commissures.
Supplementary Table 2: Material properties of valve constituents and biomaterials.
*The measured values in parenthesis were corrected to 0.53 mm for stability of the finite
element simulation. The material constants were obtained based on the corrected
thicknesses.
APP: autologous porcine pericardium, HEM: HEMASHIELD, CorM: Cormatrix, BPP:
bovine pericardial patch, SJM: St. Jude Medical patch.
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Tables and Legends:
Table 1: Comparison of unrepaired (baseline controls) and repaired valve
hemodynamics for biomaterials tested with combined cardiac outputs of 2.5, 5.0 and 6.5
L/min. Median with interquartile ranges are presented.
Materials: VO (cm/s) VSC (cm/s) VC (cm/s) GOA (cm2) LVW (mJ)
Control 157 (140-213) 6.3 (4.5-8) 68 (53-76) 5.74 (5.29-6.87) 1002 (703-1238)
APP 113 (67-163) 5.35 (2.6-6.9) 48 (44-55) 5.25 (4.16-5.99) 1016 (932-1229)
P-value 0.03 0.2 0.02 0.08 0.6
Control 136 (109-167) 3.3 (0.3-7.5) 63 (59-71) 5.91 (5.58-6.43) 1018 (689-1234)
HEM 120 (100-164) 4.4 (2.4-6.7) 48 (35-61) 5.06 (4.24-5.73) 1106 (763-1249)
P-value 0.3 0.4 0.01 0.05 0.3
Control 219 (188-241) 6.2 (5.2-13.1) 81 (69-86) 6.08 (5.82-6.83) 1033 (694-1221)
CorM 111 (76-150) 4.3 (1.5-6) 52 (44-63) 5.27 (4.57-5.56) 1086 (727-1241)
P-value 0.0001 0.01 0.0001 0.0001 0.4
Control 177 (153-213) 7.25 (4.7-11.3) 65.5 (58-83) 5.45 (4.99-6.36) 1039 (709-1222)
BPP 173.5 (127-195) 4.05 (3.3-7.6) 72.5 (57-80) 5.08 (4.54-6.59) 1059 (730-1253)
P-value 0.4 0.08 0.8 0.5 0.4
APP (N=21): autologous porcine pericardium, HEM (N=21): HEMASHIELD, CorM
(N=15): Cormatrix, BPP (N=18): bovine pericardial patch, VO: valve opening velocity,
VSC: valve slow closing velocity, VC: valve closing velocity, GOA: maximum geometric
orifice area, LVW: left ventricular work.
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Table 2: Valve hemodynamics in un-repaired porcine valves at different cardiac
outputs (CO).
Measure Combined (N=75)
CO 2.5 L/min (N=25)
CO 5.0 L/min (N=25)
CO 6.5 L/min (N=25)
P-value
VO (cm/s) 171 (140-213) 145 (116-159) 181 (142-219) 194 (164-220) 0.002
VSC (cm/s) 5.9 (2.9-8.8) 7.5 (5.5-9) 5.8 (1.3-9.5) 5.2 (2.2-7.2) 0.12
VC (cm/s) 68 (59-79) 61 (48-73) 68 (60-83) 76 (63-83) 0.05
GOA (cm2) 5.82 (5.37-6.5) 5.8 (5.4-5.99) 6.06 (5.32-6.56) 5.83 (5.31-6.67) 0.5
LVW (mJ) 1029 (703-1222) 686 (661-703) 1032 (1000-1051) 1249 (1229-1283) 0.001
VO: valve opening velocity, VSC: valve slow closing velocity, VC: valve closing
velocity, GOA: maximum geometric orifice area, and LVW: left ventricular work.
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Table 3: Summary of FEM simulation results..
Measure Original APP HEM CorM BPP SJM
VO (cm/s) 70 65 -7% 64 -9% 66 -6% 65 -7% 66 -6%
VC (cm/s) 55 53 -4% 52 -5% 53 -4% 55 0% 53 -4%
GOA (cm2) 5.6 5.2 -7% 5.1 -9% 5.3 -6% 5.2 -7% 5.3 -6%
NCC (kPa) 1,746 1,841 +5% 2,263 +30% 1,973 +13% 1,946 +12% 1,809 +4%
L/R cusp (kPa) 1,746 2,206 +26% 1,895 +9% 2,309 +32% 2,212 +27% 2,159 +24%
Combined Stress (kPa)
3,492 4,047 +31% 4,158 +39% 4,282 +45% 4,158 +39% 3,968 +28%
NCC: maximum Von Mises stress in the replaced non-coronary cusp; L/R cusp:
maximum Von Mises stress in the left and right non-replaced coronary cusps, APP:
autologous porcine pericardium, HEM: HEMASHIELD, CorM: Cormatrix, BPP: bovine
pericardial patch, VO: valve opening velocity, VC: valve closing velocity, GOA: maximum
geometric orifice area
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Supplementary Material (2 Tables):
Supplementary Table 1: Unpressurized porcine aortic valve dimensions (mm).
LVOT STJ H RSM LH FE Hs Hc
24 22 17.8 15 17 34 19 9.6
LVOT: diameter of left-ventricular outflow tract, STJ: diameter of sinotubular junction, H:
valve height, RSM: maximum radius of aortic sinuses, LH: leaflet height, FE: leaflet free
margin, Hs: sinus height, Hc: height of commissures.
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Supplementary Table 2: Material properties of valve constituents and biomaterials.
*The measured values in parenthesis were corrected to 0.53 mm for stability of the finite
element simulation. The material constants were obtained based on the corrected
thicknesses.
Material c1 (kPa) c2 (kPa) c3 (kPa) Thickness (mm)
Porcine aorta 6.35 2.62 3.72 1.9 Porcine leaflet 2.01 221 48.2 0.53
APP 3.00 11,200 375 0.53 (0.12)* HEM 2.26 2,130 1,060 0.74 CorM 3.00 15,700 111 0.53 (0.19)* BPP 2.59 15,300 376 0.48 SJM 3.00 4,180 186 0.53 (0.21)*
APP: autologous porcine pericardium, HEM: HEMASHIELD, CorM: Cormatrix, BPP:
bovine pericardial patch, SJM: St. Jude Medical patch.
