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Chapter 28
https://t.me/med1917
Experimental Designs for In Vitro
Assessment of Valve Thrombosis
Ali N. Azadani1and Danny Dvir
1
University of Denver, Denver, CO, United States;2University of Washington, Seattle, WA, United States
2
INTRODUCTION
Valvular heart disease is a major cause of cardiovascular morbidity and mortality worldwide. Each year, more
than 100,000 prosthetic heart valve replacements are performed in North America, and over 300,000 are performed
worldwide [1e3]. The numbers are expected to triple by 2050 owing to aging populations and the relatively
high incidence of rheumatic heart disease in developing countries [4e7]. Two ty pes of prostheti c heart valves are often
used in the clinical environme nt: mecha nica l and bioprosthetic heart valves. Mechanic al valves are more durable
than bioprosthetic valves, but they are more thrombogenic. Traditionally, the risks of lifelong systemic anticoagulation
therapy following the surgical implantation of a mechanical heart valve were evaluated versus the risks of
bioprosthetic valve deterioration and possible reoperation. Owing to promising improvements in the long-term durability
of surgical bioprosthe se s, bioprosthetic heart valves have been implanted in the m ajo rit y of patients who are over
60 years of age to avoid life lo ng anticoagulation therapy [8]. Moreover, transcatheter aortic valve replacement (TAVR)
hasemergedasasafeandeffectivealternativeforpatientsdeemed at high/intermediate risk for surgical aortic valve
replacement (SAVR) [9e11]. As a result, the possibility of redo t ranscatheter aortic valve (TAV)-in-valve implantation
also contributes to the sh ift toward the use of bioprosthes es in patients with severe sympt om ati c aortic stenosis [12e14].
Therefore, the majority of patients who undergo valve replacement currently receive bioprosthetic heart valves
[8,15e17].
Although heart valve replacement is an ef fe ctiv e and safe treatment, valve thrombosis is a serious and important
complication of heart valve replacement. Owing to the inherent t hrom bo ge nici ty of mechanical heart valves, long-term
anticoagulation therapy is required in all patients who receive mechanical heart valves to prevent thrombotic complications [17]. In contrast, bioprosthetic valves are generally believed to be less thrombogenic than mechanical valves.
However, there is an increasing awareness of decreased leaflet mobility of bioprosthetic heart valves [18e21].Although
often subclinical, leaflet thrombosis may potentially lead to an increased risk of stroke and transient ischemic attacks
[18]. An u pda ted overview of prosthetic valve th romb o sis and its thromboembolic complications can be found in
Chapter 27, “Prosthetic Valve Thrombosis: Biotissue Valves, Mechanical Valves, and TAVR Valves.” In addition to
pharmacological preventio n methods, it is important to investigate the predisposing factors affecting prosthetic valve
thrombosis. Minimizing the thrombogenicit y of heart valves is essential to ensure long-term undisturbed valve function.
Therefore, improvements in valve design and materials have been investigated to reduce valve thrombogenicity and
improve outcomes in patients who undergo heart valve replacement.
Despite significant advances in our understanding of thrombosis over the past 150 years, Virchow’s triad has been
considered as an explanation for the etiology of clot formation [22]. Rudolf Virchow proposed three factors that predispose
patients to thrombosis: (1) abnormal changes in the vessel wall, (2) abnormal blood flow, and (3) abnormal changes in
blood constituents [23]. Extensive changes in these variables are evident in prosthetic heart valve replacement. Abnormal
changes in the vessel wall represent the biocompatibility and surface properties of the prosthesis itself (material and design)
as well as host tissue damage. Hemodynamic factors include flow characteristics of the prosthetic valve, site of implantation, and overall cardiace
hemodynamic status of the patient. Hemostatic factors involve characteristics of the blood
Cardiovascular Thrombus. https://doi.org/10.1016/B978-0-12-812615-8.00028-4
Copyright © 2018 Elsevier Inc. All rights reserved.
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constituents of the patient (hypercoagulability) as well as adequacy and duration of antithrombotic treatment. Clinically,
the early postoperative/postintervention period represents a challenge, with the need to balanc e the risks of excessive
antithrombotic therapy and associated bleeding complications in patients undergoing prosthetic heart valve replacement. In
the following, we will review the most relevant contributions in the literature to in vitro assessment of prosthetic heart
valve thrombosis and will discuss attempts toward minimizing the risk of valve thrombosis.
BLOOD COMPATIBILITY ASSESSMENT OF MATERIALS
Material selection is an important consideration in design and development of blood-contacting medical devices.
The selection encompasses important considerations including type of blood contact (direct or indirect), duration of
blood exposure, preferred surface chemistry and topology, and physical and mechanical properties of the materials to reach
the desired function of the device. Blood compatibility is defined by the extent to which the coagulation and inflammatory
processes are destructive to the host or to the medical device [24]. The most common failure mechanism of implanted
medical devices that arise from the lack of hemocompatibility is associated with thromboembolic complications.
Upon direct blood contact with (foreign) materials utilized in medical devices, the blood coagulation system becomes
activated and the process can lead to device thrombosis [25]. Thrombosis is a multicomponent pathology that can be
triggered by several different platelet aggregation pathways and may involve the intrins ic and extrinsic coagulation
pathways [26]. Other complications include hemolysis, infection, and structural and nonstructural dysfunction. Here,
we explore the current understanding of these interactions and discuss practical testing requirements with a focus on
hemocompatibility testing of prosthetic heart valves.
Early materials developed for use inside the human body such as silk, gold, silicone, rubber, stainless steel, and
titanium had a common feature of biological inertness to reduce the immune response to the f oreign body [27].
Significant progress has been made in the past few decades in the field of biomaterials [28]. Howev er, the materials used
in the fabrication of medical devices are typically chosen from a relatively short list of materials that have been highly
testedandpreviouslyapprovedforuseinhumans[24]. The approach is driven by the fact that biological evaluation of
new noneclinically tested materials can be costly, demanding, and potentially risky. Although some materials such as
silicones and metals have a poor hemocompatibility, they are s till commonly used based on their history and the need for
their mechanical properties. The commercially available prosthetic heart valves feature materials such as metals
(e.g., stainless steel, cobalt alloys, titaniumeni ckel alloy [nitinol], titanium, and titanium alloys), ceramics
(e.g., polycarbonate), polymers (e.g., Dacron, Teflon, polyester , polytetr afluoroethylene, polyurethane), and biological
tissue (e.g., bovine, porcine, and cadaveric tissue).
Important initiatives and developments have emerged to improve the hemocompatibility of materials using
surface modification [29]. Strategies to improve biomaterial hemocompatibility have two main purposes: (1) reduction of
blood-formed elements and protein interaction and absorption on the surface, e.g., using bioinert coatings and hydrogels to
reduce initiation of blood coagulation cascade reactions, and (2) limiting the effect of the antithrombogenic drug near the
medical device without systemic suppression of hemostatic function. As an example, heparin conjugation to biomaterials,
due to its potent anticoagulant properties, was widely used to reduce the thrombogenicity of biomaterials in contact with
blood [30]. Several variations of surface-immobilized heparin have been developed and applied to a wide range of medical
devices, including stents [31], vascular grafts [32], and catheters [33]. Nevertheless, the actual clinical performance of
blood-contacting devices with a heparin coating has been varied [34]. It may be due to the fact that heparin is only an
anticoagulant and not an antiplatelet agent [34]. In certain devices where platelets play a major role in thrombus
development, such as left-ventricular assist devices, hemodialysis catheters, and extracorporeal devices, heparin coatings
have failed to demonstrate consistent antithrombogenic efficacy. Consequently, surface-bound heparin coating has not
eliminated the need for systemic antithrombogenic drug usage in most cases [25]. Recognizing the significance of
incorporating antiplatelet properties onto biomaterial surfaces, nitric oxide-coated surfaces [35,36], glycoprotein IIb/IIIa
inhibitors [37], and in vitro seeding of endothelial cell types onto biomaterials [38] have also been considered for device
coating. However, none of the antiplatelet biomaterial surfaces have been in routine clinical use. In the long run, tissue
engineering has the potential to significantly improve the hemocompatibility of implanted devices by the regeneration of
blood-contacting functional tissue. Tissue-engineered heart valves in particular would adapt with the patient and have the
potential to improve the quality of life of millions of patients with valvular heart disease [39e41].
Hemocompatibility and evaluating interactions of medical devices with blood are required by regulatory agencies.
Hemocompatibility tests are described in International Organization for Standardization (ISO) standa rd 1099 3-4: 20 02
and its amendment ISO 10993-4:2006. The standard has been under revision and the revised standard was published in
April 2017. The standard provides general requirements for assessing the interactions of medical devices with blood and
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suggests a process diagram to aid in deciding when new testing is required as opposed to when supporting risk
assessment and historical data may be utilized. The standard does not offer specific methodologies and study techniques,
but emphasizes important consideration and strategies for preclinical testing. The potential types of deviceeblood
interaction in the 2017 revision of the ISO standard (ISO 10993-4:2017) are classified into two main categories: (1)
hemolysis and (2) thrombosis. Hemolysis i s split into two separate categories: material-induced and mechanically
induced hemolysis. Thrombosis, on the other hand, is explicitlydescribedtobeaninvivo/exvivophenomenon,but
coagulation, platelet activation , comple ment , and hematology can be simulated using in vitro tech niqu es. M ech anic al
and bioprosthetic heart valves are implanted devices with constant interaction with v ario us blood c om ponen ts. As a
result, the standard requires testing for hemolysis and thrombosis for p rosthetic heart valves. It is essential that
the preclinical testing simulate clinical conditions such as flow conditi on s, duration of contact, temperature, sterile
condition, and anticoagulation as much as possible. For mechanical heart valves, it is recommended to evaluate both
material-induced and mechanically induced hemolysis. However, for tissue hear t val ve s on ly mat er ia l-ind uce d hem olysis is recommended to be evaluated by ISO 10993-4:2017.
In Vitro Hemocompatibility Testing of Biomaterials
To reduce the use of animals in safety and risk assessment studies for hemocompatibility of materials and medical
devices, there is a constant need for verification and validation o f alternative testing methodologies such as in vitro
testing systems. In addition, the composition of human blood is considerably different from th at o f o ther species [42,43].
Therefore, in vitro circulation models are at tractive and promising tools to use with human blood and to test the
hemocompatibility of materials/devices. In general, in vitro methodologies are useful in evaluating basic hemocompatibility properties of materials/devices under controlled conditions in isolation from confounding factors such as
tissue injury associated with in vivo usage. Furthermore, in vitro tests allow repeating the assays for statistical purposes
and permit testing again st reference m at er ials (controls) using the same batch of blood. Nevertheless, in vitro tests may
not be accurate to predict blood interactions with materials/devices with prolonged or permanent contact because of
blood stability issues in the in vitro systems.
The in vitro models serve as the worst case scenario because activation products accumulate due to the continued
contact and absence of component removal by organs such as liver and kidneys. In addition, endothelial cells, with their
capacity to inhibit blood coagulation, are absent in the in vitro models [44]. Static in vitro experiments provide some
limited information that can be utilized to screen biomaterials and determine platelet adhesion/spreading on the test
materials [45,46]. However, materials and devices in clinical applications contact blood under flow conditions. As a
result, dynamic evaluation o f biomaterialsisessentialtoassessthehemocompatibili ty of materials/devices. In the
dynamic systems, blood temperature, duration of contact, and flow rheological conditions such as flow rate, wall shear
stress, and pressure must be as similar as possible to those in the clinical setting to achieve pertinent results. In particular,
shear stres s and duration of its exposure are key determinants of blood activation and cell adhesion [47,48].Ithasbeen
shown that high shear stress forces can activate platelets and may also cause hemolysis in prosthetic heart valves [49,50].
In vitro flow systems s uch as par al lel -pla te chambers [51,52],viscometers[53,54], tubular systems [55,56],and
miniature pulse duplicator systems have been used to study hemocompatibility aspects (thrombosis and hemolysis) of
biomaterials and medical devices such as prosthetic heart val ves . To evaluate mech anic al ly induce d hemolysi s and
platelet activation, in vitro testing has to be performed at the highest blood flow rate observed in the clinical flow range.
On the other hand, for thrombosis, the minimum labeled blood flow rate may be essential to characterize the safety of the
devices.
In the in vitro tests, it is important to specify blood volume in t he system, blood temperature, flow condition,
anticoagulant type, and blood contacting surface area of the test system. Moreover, since coagulation protein activation
is proportional to blood-contacting surface area, the test sample surface area to blood volume ratio (exposure ratio)
should be specified in each study. Exposure ratios of 3.0e6.0 cm
with ISO 10993-12. In addition, blood collection and preparation a re vital steps to succeed in hemocom patibility testing.
However, as of this writing, there is no agreement on which prean alyt ic al test should be performed to characterize
the functionality and prove the suitabil ity of the donated blood. Nevertheless, hemocompatibility testing should be
performed using blo od obtained from health y nonsmoker donors, free of medication such as aspirin or other platelet
function inhibitors [45]. The average platelet activation in healthy donors who have received no medication is about 6%
[57]. All the hemocompatibility tests should be performed within 4 h of blood collection. It has been clearly shown
that platelet function and white blood cell activity are significantly influencedafter4hofstationary sto rag e at room
temperature [58].
2
/mL blood (based on device thickness) are consistent
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It is also imperative to simulate clinical anticoagulation conditions in the in vitro tests as much as possible. As a result, a
commonly used clinical concentration of heparin (1.5 IU/mL) is often used in the in vitro tests [44]. Furthermore, donor
specific heparin concentration can be determined by static pre-tests to reduce the inherent variability in coagulability
between donors [174]. In addition to heparin, other anticoagulant (e.g., sodium citrate and hirudin for studies with focus on
plateletebiomaterial interactions, and ethylenediaminetetraacetic acid) as well as antiplatelet agents have been utilized in
in vitro studies [45]. Each one of the agents has unique advantages and disadvantages in verifying hemocompatibility of
biomaterials [44,59]. The choice of agents should be driven by the clinical background and the scientific inquiries of the
study. Nevertheless, there is no accepted guideline that defines the appropriate concentration of each of the anticoagulants
utilized in in vitro tests, which prohibits direct comparison o f different materials and studies.
Parallel-plate flow chambers are widely used to study platelet activation and thrombus formation. They enable a surface
to be precoated with thrombogenic materials to measure platelet adhesion, platelet aggregation, and coagulation within one
experiment. The device is easy to sterilize and allows whole blood perfusion through the chamber at well-defined shear
conditions. Platelet adhesion, activation, and aggregation can be measured by epifluorescence/radiolabeling videomicroscopy, flow cytometry, and particle counter, respectively. An overview of substrates available for coating and a
practical guide for the use of parallel-plate flow chambers have been given by Kruchten et al. [51]. In addition to the
parallel-plate flow chambers, parallel-plate viscometers and cone-and-plate viscometers have been used to investigate
plateletebiomaterial interactions [54]. Platelet adhesion, activation, and aggregation can be measured over a range of shear
rates in viscometers. In comparison with the parallel-plate flow chambers, viscometers offer some unique advantages in
studying the blood-contacting characteristics of a biomaterial, such as (1) the ability to generate uniform and constant shear
rate flow fields in the viscometer; (2) eliminating the use of a pump to circulate the fluid, which may potentially cause
blood damage; and (3) removing all of the connecting tubes in the recirculation system to which the circulating platelets are
exposed prior to the test surface.
Tabular systems such as the Chandler loop have also been used extensively to test the hemocompatibility of biomaterials [60e63]. The classical Chandler loop system consists of a closed tubing partially filled with blood that circulates
constantly using a step motor [55,64]. Blood cascades will be activated in this approach, but protein denaturation is also
common at the bloodeair interface [65,66]. It has clearly been demonstrated that bloodegas interfaces activate blood
cascades [67,68]. As a result, modified versions of the Chandler loop system such as the Haemobile [69,70] and small
roller pumps [71,72] have been used in the past by investigators to avoid bloodeair interfaces. In the tabular systems, it is
important to simulate clinical flow conditions such as wall shear stress and flow rate as close as possible. It has been shown
that blood flow pulsatility significantly in
fluences platelet adhesion on biomaterials [73].
Since fluid dynamics and cellular-level mechanisms act synergistically in thrombus formation in prosthetic heart valves,
a few studies have attempted to study hematology and fluid dynamics using miniature pulse duplicator systems. The
thrombosis tester of the Helmholtz-Institute Aachen was one of the first miniature pulse duplicators that were designed and
utilized to determine the thrombogenic potential of heart valves [74,75]. The pulse duplicator mimics the physiological
pressure and flow conditions to a reasonable level. Human and porcine blood have been used in the in vitro analyses. More
recently, Arjunon et al. [76] used a custom-designed pulse duplicator to determine the thrombogenic potential of prosthetic
heart valves. The system provides a hemocompatible small-volume test platform that can be used for thrombogen icity
studies. In addition, particle image velocimetry was utilized in the miniature pulse duplicator by using a blood-mimicking
fluid to characterize the flow patterns in the system. These in vitro tests enabled design optimization of prosthetic heart
valves and provided a suitable supplement to animal studies.
In all in vitro tests, it is recommended to determine the cell count (erythrocytes, platelets, leukocytes), platelet
function analysis, and blood gas a nd pH data before starting the experiments to confirm blood quality and determine the
baseline characteristics. In hemocompatibility tests, commercially available enzyme-linked immunosorbent assay kits
and flow-cytometric assay can be used to charac ter ize th e bloo d reactions. For platelet activation, parameters such as
platelet count (% loss), b-thromboglobulin, and platelet factor 4 can be used as measures of platelet activation
[56,58,63]. In addition, since thromboxane A
is highly unstable and rapidly converts to thromboxane B2(TXB2), TXB
2
can be used as an indicator to monitor platelet activation [77]. In addition, several studies point to the im por ta nce of
quantifying platelet-leukocyte aggregation as a sensiti ve measure of platelet activation [175e 179]. For blood
coagulation cascade, the conversion of prothrombin to thrombin is an important reaction. Furthermore, antithrombin III
is an important coagulation inhibitor that neutralizes thrombin by forming the thrombineantithrombin (TAT) complex.
Thus, TAT plasma concentration measurements can serve as an indirect marker for the detection of activation of the
coagulation cascade in in vitro systems [63]. Other coagulation paramet er s such as partial thromboplastin time, F1.2
(protein fragment released from prothrombin upon formation of thrombin), and FPA (protein fragment released from
fibrinogen upon formation of fibrin) can also be determined in in vitro tests. For hemolysis, photometric and colorimetric
2
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tests can be employed to determine the amount of free hemoglobin in plasm a. More ov er, C3a an d SC5b -9 can be
used to characterize the complemen t system, and polymorphonuclear elastase released by the stimulation of leukocytes
can b e used as an indicator of leukocyte activation and inflamm ation in hemocompatibility test s. In addition, adhesion
of blood cells such as platelets on the sample surface and changes in morphology can be examined by light and scanningelectron microscopy. Seyfert et al. [78] reported a comprehensive test collection for platelet activation, hemolysis,
thrombin generation, complement activation, contact activation, fibrinogenefibrin conversion, fibrinolysis, and
proteolysis.
A variety of negative and positive reference materials have been used in in vitro systems. ISO 10993-4 recommends
using blood plasma itself without the material/device as a negative control. In addition, ISO 10993/12 specifies
polyvinylchloride (PVC) and polyurethane as negative reference materials. The most commonly used negative refere nce
materials are polyethylene and PVC and the commonly used positive materials are latex rubber and glass. Although ISO
10993-4 presents the general requirements to characterize blood interactions, a major shortcoming is the lack of acceptance
criteria. New materials and devices can be compared with well-characterized biomaterials or relevant existing devices
already in clinical use, and can be accepted when hemocompatibility results show similar or better results. Nevertheless,
since the designs of the in vitro hemocompatibility testing systems are often different and not always well known, any
comparison of the data and conclusions from different systems should be considered with care.
In addition to ISO 10993-4, the ASTM F2888-13 test method assists in the evaluation of medical devices for
their ability to induce thrombus formation. The in vitro test will determine if biomaterials exposed to human whole
blood will adversely affect the platelet and leukocyte counts in whole blood. This assay may be part of the
hemocompatibility evaluation for devices and materials contacting human blood, as per ANSI/AAMI/ISO 10993-4.
Moreover, ASTM F756-00 provides a standard protocol for the assessment of hemolytic properties of biomaterials used in
the fabrication of medical devices. The protocol evaluates the acute in vitro hemolytic properties of materials in contact
with blood under static conditions. Finally, as discussed above, the in vitro blood flow models provide a valuable
evaluation tool to assess medical device thrombosis. However, the accuracy of the in vitro tests may not be satisfactory for
bloodedevice interactions occurring upon prolonged and permanent contact. As a result, evaluations of the devices are
frequently necessary in in vivo animal models to assess the blood compatibility of biomaterials.
In Vivo Animal Models for Hemocompatibility Testing
Considering the limitation of in vitro models to assess hemocompatibility and thrombogenicity, animal studies remain an
important step in the evaluation of biomaterials and medical devices [79]. To date, many animal models have been
developed to evaluate the safety of prosthetic heart valves, including canine, porcine, calf, baboon, and adult ovine models
[42]. Traditionally, because of the large amount of information and data available in the surgical literature, the canine
model was considered to be the gold standard for heart valve replacement. In addition, the canine model was considered to
be more prone to platelet adhesion [80,81], thrombosis [82], and hemolysis [80]. As a result, the canine model was viewed
as an advantage when evaluating valve thrombosis in preclinical studies [83,84]. However, because of ethical concerns,
there has been a shift from the canine model to farm animals. Pig hearts are similar to human hearts in regard to structure,
size, and function [85]. Moreo ver, their coagulation cascade is quite similar and comparable to that of humans. Despite the
similarities, several problems have been recognized in using the swine model, including significant bleeding complications
due to difficulty in maintaining safe levels of anticoagulation therapy and the rapid growth of young swine during
the follow-up period [86]. At this writing, the most widely used animal model for evaluating prosthetic heart valves is the
ovine model [84]. Normal cardiovascular physiological parameters of the ovine and its coagulation profile are similar to
those of humans [87]. As a result, they are the primary model used to assess hemocompatibility and thrombogenicity of
prosthetic heart valves.
According to ISO 10993-4, ISO 5840-2, and ISO 5840-3 standards, thrombosis and hemolysis testing is required for
both surgically implanted heart valve substitutes and heart valve substitutes implanted by transcatheter techniques.
Pressur e and flow measurements are important in determining prosthetic heart valve performance [88].Forhemolysis
assessment, red blood cell count, hematocrit, reticulocyte count, lactate dehydrogenase, haptoglobin, and plasma-free
hemoglobin should be assessed at minimum. Addition a l hematology and clinical chemistry analyses should be
conductedtoassessinflammatory response, platelet consumption, and liver and renal function. In addition, a full
postmortem exam is recommended to d isc lose peripheral thromboemboli both macroscopically and microscopically.
Furthermore, a histolog ic description of the thrombotic materia l should be provided. Moreover, au tops y of the valve and
adjacenttissuesisrecommended.Itisimportant to note that the thrombotic and inflammatory response assessed in
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animal models is not completely equivalent to human blood conditions because of the differences in blood reactivity
between species and species differences in blood reactions to medicat ions. As a result, the interpretation of the results in
the animal models must be done carefully consider ing that results may be different in h um ans.
HEMODYNAMIC FACTORS
Another aspect of Virchow’s triad is abnormal blood flow. Fluid mechanics and hematology act synergistically to induce
thrombotic complications [89,90]. Valve thrombosis and hemolysis may be initiated by nonphysiological hydrodynamic
forces in addition to bloodstream exposure to foreign surfaces. Elevated shear stress in the blood flow can activate platelets
and damage red blood cells [91]. In the case of platelets, an increase in shear-induced deformation of the platelet membrane
could promote platelet activation through a series of morphological changes in platelets that involve reorganization of the
actin and microtubule components of the cytoskeleton [92,93]. In red blood cells, an increase in shear-induced deformation
beyond the erythrocyte membrane threshold may lead to membrane rupture and instant lysis or permanent deformation and
subsequently cell fragmentation [94]. It has been shown that the true shear force experienced by platelets and red blood
cells is created by viscous shear stresses rather than Reynolds shear stresses, when turbulence is present in the blood flow
[92,93].
Flow-induced blood cell damage has been studied in different flow systems such as viscometers, jet flows, and
parallel-plate flow cham be rs [95e99]. These studies have established a quantitative link between blood cell damage and
the mechanical for ce environment experienced by bl oo d elements. Activation of blood elements is related to she ar stress
magnitude as well as exposure time. The time of exposure is an important variable and the higher it gets, the lower will
be the stress threshold level of blood cell damage. The shear stress threshold under which no red blood cell damage has
been detected is 150 N/m
2
10
s, while Bacher and Williams [101] reported a threshold level of 500 N/m2for an exposure time of 102s. On the
other hand, Rooney et al. [102] found a thr e shold level of 450 N/ m
blood cells, platelet activation o ccurs at significantly lower shear stress thresholds. Hung et al. [103] reported platelet
damage at 10e16. 5 N/ m
time of 10
30e100 N/m
3
s. In addition, Ramstack et al. [105] reported platelet activation at a somewhat higher threshold of
2
with exposure time of 102e101s. In addition to the experimental findings, several mathematical models
have been proposed for shear-induced platelet adhesion, activation, and aggregation as well as red blood cell damage
[106e109].
In prosthetic heart valves, fluid dynamics has been known to play a critical role in thrombotic and hemolytic
complications [89,110]. The thromboembolic events due to platelet activation and hemolysis, caused by rupture of
erythrocyte membranes, have been correlated to high shear stress regions in the blood flow [90,111,112]. Numerous
researchers have identified regions of high shear stress in prosthetic heart valves using experimental testing [113e119]
and computational modeling [107,120]. M echanical heart valves would serve as a good example. Throm bu s form ation i n
explanted bileaflet mechanical valves was commonly observed within the hinge region [121,122]. In vitro studies and
experimental methods have been instrumental in coupling the fluid dynamics o f prosth etic he art va lve s to throm bo tic
complications. Experimental fluid d yn am ics techniques such as laser Doppler velocimetry and particle image velocimetry have been extensively used to characterize flow through mechanical heart valves under steady and pulsatile
conditions [123]. Alongside experimental studies, computational fluid dynamics is powerful simulation technique that
has been used to obtain unprecedented detail into the hemodynamics of prosthetic heart valves [49,107,120,124e126].It
was found that bileaflet mechanical heart valves are associated with high shear stress levels, long exposure times to the
shear stresses, and flow recirculation at the hinge regions [50]. As a result, various design modifications of the hinge
region were explored to reduce the level of thromboembolic events [119,127].
In addition to high shear st ress flow regions, regions of blood flow stagnation and stasis promote elevated transport
of blood c om pone nt s to the biomaterial surface, increase interplatelet collisions, and provide an o p port unity for
platelets and blood proteins to accumulate to critical concentrations, leading to thrombosis [25,106,128e131].Asa
result, prolonged blood residence time (BRT) and stasis may be part of the mechanism behind the formation of
leaflet thrombosis in surgical and transcath eter bioprosthetic heart valves. Clinical observation shows that the risk of
valve thrombosis is higher for right-sided prosthetic heart valves than for left-sided heart valves, e.g., obstruction of a
tricuspid mechanical prosthesis due to valve thrombosisis20timesmorefrequentthaninleft-sidedmechanical
heart valves [132]. In addition, the risk of mitral valve thrombosis is 2 e 3 times higher than that of a n aortic
prosthesis [132].
2
[100]. Leverett e t al. [100] reported a threshold level of 150 N/m2for an exposure time of
2
for an exposure time of 103s. In comparison to red
2
and exposure time of 102s, while Williams [104] found a threshold of 13 N/m2and exposure
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Bioprosthetic heart valves are generally believed to be less thrombogenic than mechanical valves. However, there is an
increasing awareness of decreased leaflet mobility of bioprosthetic heart valves. Valve thrombosis and leaflet immobility in
bioprosthetic valves are detectable only by four-dimensional computed tomographic or transesophageal echocardiographic
imaging [133]. The thrombotic materials in bioprosthetic heart valves are predominantly found on the aortic side of the
leaflets [133e135]. It has been shown that the rate of subclinical leaflet thrombosis in bioprosthetic aortic valves is
significantly more common in TAVs than surgical aortic valves (SAVs) [21,133e135]. However, the underlying mechanisms that facilitate thrombus formation on leaflets, particularly following TAVR and valve-in-valve (ViV) procedures,
are not well addressed.
In 2017, in Vahidkhah et al. [136], we showed that increased blood stasis on the aortic surface of TAV leaflets may
act as a precursor to le aflet thrombosis. Diminished blood flow does not allow d ispersion of the clotting factors and
therefore the prolonged BRT on the leaflets can serve as a permissive factor in the growth of thrombi on the TAV leaflets
[26,137]. In c ontrast to SAVR, where the native calcified leaflets are removed from the annulus during the open heart
surgery, TAVs are implanted within native calcified valves in TAVR or within bioprosthetic valves in ViV procedures.
As a result, unlike surgical bioprostheses, the TAV frame may be confinedbytheleaflets of the calcified native valve in
TAVR (Fig. 28.1, left) or by the leaflets and the frame of the degenerated bioprosthetic valve in the ViV setting
(Fig. 28.1, right). Vahidkhah et al. [138] performed a quantitative assessment of the BRT on the noncoronary and
coronary leafletsofanSAVandaTAVdevicewithanintraannulardesigninapatient-specificgeometryusing
fluidesolid interaction modeling (Fig. 28.2). Fig. 28.3 shows the contour s of BRT on the th ree le aflets of the two models
during one cardiac cycle. The top row corresponds to the SAV and the bottom row corresponds to the TAV model. In the
SAV model, the are as of high BRT were la rge st on the noncoronary leaflet. This may indicate a higher probability of
thrombus formation on the noncoronary versus the coronary leaflets following SAVR. In the TAV model, however,
similar patterns of BRT distribution were observed on all three leaflets near the fixed boundary edge (Fig. 28.3B). The
similar pat terns of BRT distribution on the three leaflets of the TAV indicate the comparable likelihood of post-TAVR
thrombosis on all three leafl et s. The results are aligned with the available clinical dat a that thrombosis has been observed
on all three leaflets following TAVR and ViV procedures [133,134,139,140]. In addition, the simulations indicated 24%,
15%, and 12% larger values of BRT in the TAV model com pare d with the SAV model on non-, right, and left coronary
leaflets, respectively. The results are also aligned wit h the c linical observ ation that leaflet thrombosis is more common in
TAVs than in SAVs [21].
Although limited in quantity, the available clinical data suggest that leaflet thrombosis is mostly observed following
the implantation of TAVs with intraannular design [20,133,141]. However, it is not clear that whether TAVs with a
supraannular design would excel in terms of the likelihood of leaflet thrombosis. As a st ep toward addres sing
this question, Vahidkhah et al. [142] investigated the effect of positioning of the TAV in the ViV setting on the flow
FIGURE 28.1 (Left) The aortic portion of a TAV stent frame is circumferentially surrounded by the calcified leaflets in TAVR. (Right) The aortic
portion of a TAV stent frame is circumferentially surrounded by the leaflets and frame of the surgical bioprosthesis in the ViV setting. TAV, transcatheter
aortic valve; TAVR, transcatheter aortic valve replacement; ViV, valve-in-valve. (Left) Reprinted with permission from Krishnaswamy A, Tuzcu EM,
Kapadia SR. Percutaneous paravalvular leak closure. Curr Treat Options Cardiovasc Med 2013;15(5):565e74.
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FIGURE 28.2 (A) Construction of the three-dimensional patient-specific geometry through processing of computed tomography angiography images.
(B) Computational three-dimensional models for (left) the surgical aortic valve (SAV) and (right) the transcatheter aortic valve (TAV). In the TAV model,
the confining geometry represents the leaflets of the native valve in TAV replacement with the intraannular function of the TAV device (bottom right) or
the leaflets and frame of the failed bioprosthesis in the valve-in-valve setting (top right). Reprinted with permission from Vahidkhah K, Javani S, Abbasi
M, Azadani PN, Tandar A, Dvir D, et al. Blood stasis on transcatheter valve leaflets and implications for valve-in-valve leaflet thrombosis. Ann Thorac
Surg 2017;104, 751e759.
dynamics, and quantified blood stasis on the leaflets in the intraannular and supraannular TAV positions. To that end,
two idealized computational models, representing ViV intraannul ar and supraannular positioning of a TAV, we r e
developed in a patient-specificgeometry(Fig. 28.4). An identical leaflet geometry was incorporated into the two models
so that the effects of TAV geometric confinement were isolated. Three-dimensional flow fields were then obtained via
the fluidestructure interaction modeling approach. As shown in Fig. 28.5, at the end of diastole, a strip of high BRT
(1.2 s) region was observed on the TAV leaflets in the ViV intraannular positioning (top r ow) at the fixed boundary
where the leaflets are attached to the frame. Such a high BRT region was absent on the TAV leaflets in the supraannular
positioning (bottom row). Thus, it was concluded that the absence of such a geometric confinement in the ViV
supraannular positioning leads to smaller BRT and subsequently less likelihood of leaflet thrombosis. In addition,
supraannular positioning of TAV devices within small surgical bioprostheses in ViV procedures is associated with
superior hemodynamics and improved leaflet kinematics [12,143e149], which may be essential for the long-term
durability of the bioprosthesis [150,151].
Additional hemodynamic factors may also affect TAV leaflet thrombosis. Despite the limited clinical data currently
available, low ejection fraction has been suggested as a predictor of leaflet thrombosis in TAVR and ViV procedures [20,21].
Experimental Designs for In Vitro Assessment of Valve Thrombosis Chapter | 28 413
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FIGURE 28.3 Snapshots showing the time evolution of the contours of blood residence time (BRT) on the valve leaflets for (A) the SAV and
(B) the TAV models during a complete cardiac cycle from the top view. BRT was significantly higher for the TAV model compared with the SAV
model. SAV, surgical aortic valve; TAV, transcatheter aortic valve. Reprinted with permission from Vahidkhah K, Javani S, Abbasi M, Azadani PN,
Tandar A, Dvir D, et al. Blood stasis on transcatheter valve leaflets and implications for valve-in-valve leaflet thrombosis. Ann Thorac Surg 2017;104,
751e759.
FIGURE 28.4 (A) Construction of the 3D patient-specific geometry via processing of computed tomographic angiography images. (B) Computational
3D idealized model for the ViV intraannular positioning of the TAV model. A sample schematic of ViV intraannular positioning is shown on the right.
(C) Computational 3D idealized model for the ViV supraannular positioning of the TAV model. A schematic of ViV supravalvular positioning is shown
on the right. TAV, transcatheter aortic valve; ViV, valve-in-valve. Reprinted with permission from Vahidkhah K, Azadani AN. Supra-annular valve-in-valve
implantation reduces blood stasis on the transcatheter aortic valve leaflets. J Biomech 2017;58, 114e122.
Vahidkhah et al. [152] demonstrated an association between reduced cardiac output and increased blood stasis on the
TAV leaflets, which can be regarded as a precursor of valve thrombosis. In addition to the low-flow regions on the TAV
leaflets, previous studies demonstrated flow stagnation in the sinuses of Valsalva following TAVR and ViV procedures
[145,153e155]. The studies agree with clinical observations on thrombosis in the sinuses of Valsalva, e.g., early
post-TAVR thrombus formation in the noncoronary cusp following CoreValve implantation [156].Finally,itisimportant
to note that, in addition to computational fluid dynamic simulations, multiscale computational m odels for platelet
activation and coagulation reactions have been developed since the late 1990s [157e159]. The computational models for
studying thrombus development supplement knowledge gained from in vitro experimental testing and clinical research.
An updat ed overview of the mathematical models can be found i n Chapter 5, “Mathematical Modeling of Thrombus
Formation Under Flow.”
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FIGURE 28.5 Snapshots showing the time evolution of the contours of blood residence time on the aortic side of valve leaflets for (A) the intraannular
and (B) the supraannular positioning of the TAV model during a complete cardiac cycle. TAV, transcatheter aortic valve. Reprinted with permission from
Vahidkhah K, Azadani AN. Supra-annular valve-in-valve implantation reduces blood stasis on the transcatheter aortic valve leaflets. J Biomech 2017;58,
114e122.
ANTITHROMBOTIC TREATMENT
The third element of Virchow’s triad is abnormal changes in blood constituents. Hemostatic factors following prosthetic
heart valve replacement involve adequacy and duration of a ntithrombotic treatment as well as the degree of hypercoagulability of patients. The early postoperative/postintervention period represents a challenge, with the need to balance
the risks of excessive antith rombotic therapy and associated bleeding complications. Patients with mechanical heart
valves receive lifelong oral anticoagulant therapy with vitamin K antagonists (VKAs), such as warfarin, to prevent
thromboembolic complications [160,161]. Neverthele ss, it has been a challenge to balance the risks of underanticoagulation against those of excessive anticoagulation [162]. Furth ermore, there is a lack of consensus regarding
the optimal intensity and timing of anticoagulation to prevent early thromboembolism after mechanical valve
replacement [163].
Since 2008, bioprosthetic heart valves have gained momentum over mechanical prosthetic valves, as discussed previously, because of the lack of long-term requirement for anticoagulation therapy, favorable clinical results, and advances
in TAVR and ViV procedures [164,165]. In general, patients with bioprosthetic heart valves are at a higher risk of ischemic
stroke or peripheral embolism than the normal population. As a result, VKAs are recommended for 90 days for mitral,
tricuspid, and pulmona ry valve implantation according to the American Heart Association/American College of Cardiology (AHA/ACC) and European Society of Cardiology (ES C) guidelines [166,167]. For aortic bioprosthetic valves,
however, VKA is recommended for 3e6 months according to the AHA/ACC guideline, while aspirin is preferred over
VKA in the ESC guidelines. Evidently, the guidelines are not concordant in their recommendations because they are based
on different observational retrospective data [168e171]. Following TAVR, dual antiplatelet therapy with clopidogrel and
aspirin is currently recommended and used in most centers worldwide, but the duration of clopidogrel varies among
studies, ranging from 1 to 6 months [172]. As of this writing, there are no clinically tested guidelines to recommend
appropriate therapy for post-TAVR thrombus formation. Most TAVR patients with valve thrombosis have been successfully managed with oral anticoagulation therapy, with significant hemodynamic improvement and resolution of
thrombus [21,133]. Nevertheless, the duration of oral anticoagulation therapy is not known and should be determined on a
case-by-case basis considering bleeding risks.
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