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In vitro evaluation of the effects of intraluminal thrombus on abdominal aortic aneurysm wall dynamics
Medical Engineering & Physics 33 (2011) 957–966
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In vitro evaluation of the effects of intraluminal thrombus on abdominal aortic aneurysm wall dynamics Florentina Ene ∗ , Carine Gachon, Patrick Delassus, Ronan Carroll, Florian Stefanov, Padraig O’Flynn, Liam Morris Department of Mechanical Engineering, Galway Medical Technologies Centre (GMedTech), Galway Mayo Institute of Technology (GMIT), Dublin Road, Galway, Ireland
a r t i c l e
i n f o
Article history: Received 20 January 2010 Received in revised form 12 March 2011 Accepted 14 March 2011 Keywords: Abdominal aortic aneurysm Intraluminal thrombus Silicone models Compliance Diametral strain
a b s t r a c t The optimum time to treat abdominal aortic aneurysms (AAAs) still remains an uncertain issue. The decision to intervene does not take in account the effects that wall curvature, intraluminal thrombus (ILT) properties and thickness have on rupture. The role of ILT in aneurysm dynamics and rupture has been controversial. In vitro testing of four silicone AAA models incorporating the ILT and aortic bifurcation was studied under physiological conditions. Pressures (P) and diameters (D) were analysed for models with and without ILT at different locations. The diametral strain, compliance and P/D curves were influenced by the presence, elastic stiffness and thickness of the ILT. In this case, the inclusion of ILT reduced the lumen area by 77% that resulted in a 0.5–81% reduction in compliance depending on ILT properties. With an increase in ILT stiffness from 0.05 to 0.2 MPa, the compliance was reduced by 81%. In the region of maximum diameter, there was a reduction of diametral strain and compliance except for the softer ILT which was more compliant throughout the proximal region. The shifting of the maximum diametral strain and compliance to the proximal neck was pronounced by an increase in ILT stiffness, thus creating a possible rupture site. © 2011 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction Abdominal aortic aneurysm (AAA) is one of the leading causes of death in the western world, being the 14th cause of death in the U.S. for men over the age of 55 [1,2]. An aneurysm is an area of localised and permanent dilatation of a blood vessel. Aortic aneurysms can develop anywhere along the length of the aorta, but the majority are located below the renal arteries and are called abdominal aortic aneurysms (AAAs). These AAAs can also extend into one or both of the iliac arteries. The morphology of the bulging in an AAA is patient specific and most commonly takes a fusiform shape. Aneurysms have a high risk of rupture if not detected and treated in time. The optimum time to treat AAAs still remains an uncertain issue. The majority of aneurysms (75%) contain intraluminal thrombus (ILT), however the extent of its presence varies [3]. The thrombus is a fibrin structure incorporated with blood cells, platelets, blood proteins and cellular debris [3]. Mechanical testing has proven that the ILT is a nonlinear elastic isotropic and inhomogeneous material [4]. Its exact role in AAA rupture has been controversial. Examination of ruptured AAAs has shown the presence of ILT at the location of failure [5]. It has been proposed that the presence
∗ Corresponding author. Tel.: +353 862050346. E-mail address: [email protected] (F. Ene).
of ILT induces hypoxia thus weakening the aneurismal wall and contributing to AAA rupture [40]. Conversely, other studies have proven that the ILT acts like a mechanical cushion with smaller stress in the underlying AAA wall [6], therefore possibly protecting it from rupture [7]. Computational studies of wall stress distribution have shown that, regardless of its mechanical properties, thrombus markedly reduces and redistributes the stresses in the aneurismal wall [7–12]. In vitro testing is an alternative to computational work for studying aneurysm mechanics and blood flow dynamics. Important developments have been made over the years to recreate the physiological conditions [13–17]. Compliant vascular replicas have been produced by researchers using arterial mimicking materials such as latex, silicone, polymers, gelatine and other materials [15,16,18–22]. Aneurismal models have included fusiform [15,23], idealised [16,24] and realistic with the iliac arteries [16,18]. These models however neglected to account for ILT. In a small number of studies the intraluminal thrombus has been approximated using gelatine solution [18] and dough [25], with no reported mechanical properties, inserted in dipped latex models. Other studies replicated ILT using customised polymer mixtures [21] and silicone [47]. To date, no experimental study has been done on the effect that ILT has on AAA wall dynamics, in terms of compliance and pressure/diameter variations. The objective of the current study was
1350-4533/$ – see front matter © 2011 IPEM. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.medengphy.2011.03.005
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Fig. 1. The planar idealised model of AAA. (A) The lumen of AAA with iliac arteries as reconstructed from CT scan; (B) an individual CT slice at level of aneurysm presenting the lumen region and the thrombus region from which the thrombus geometry was obtained; (C) the best fit circles of the lumen with a straight centreline; (D) the final idealised abdominal aorta with aneurysm, ILT and iliac arteries.
to analyse experimentally the effects of pulsating pressure on an idealised aneurismal aorto-iliac model with and without intraluminal thrombus. Influencing factors contributing to the complex AAA mechanics such as thrombus stiffness, thrombus distribution throughout the aneurysm sac and the aneurismal wall curvature were evaluated. The specific aims were to evaluate compliance, diametral strain and pressure/diameter variations under physiological conditions and identify the effects of ILT on wall dynamics. The results of this study can be applied to validate numerical studies in complex AAA geometries. 2. Methods
models. The wax (Texwax FG890, Texaco, Germany) created the lumen cavity. The silicones used for the ILT and arterial wall are listed in Table 1. These silicones (Wacker Chemie AG, Germany) were translucent two-component (A and B) silicone elastomers that was mixed in 10:1 ratio. Silicone fluid (Dow Corning, UK) was added in different percentages to lower the elastic modulus. For this study, four AAA models were created, one model without thrombus (M1) and three models with thrombus (M2–M4) as presented in Table 1. Three different thrombus properties were modelled, with elastic modulus of 0.05 MPa (M2), 0.1 MPa (M3) and 0.2 MPa (M4). These elastic values showing different ILT maturity are within the range (0.04–0.27 MPa) documented in the literature [4,8,26].
2.1. Physical model generation
2.2. Dynamic physiological testing of AAA models
A computed tomography (CT) scan of a male subject, suffering from an abdominal aortic aneurysm was obtained from Materialise Medical Datasets (Leuven, Belgium). The scans had a slice thickness of 5 mm with a pixel size of 0.8242 mm. A three-dimensional (3D) aortic model was created using Mimics (v11, Materialise, Leuven, Belgium) through the process of segmentation. This model exhibited non-planar curvature due to its course through the abdomen as shown in Fig. 1(A). A planar idealised model was generated by eliminating the effects of the curvature. At each scan slice, the lumen and the thrombus areas were generated (Fig. 1(B)) and a best fit circle was obtained for each area. The realistic vessel centrelines were converted to straight lines of equal length. The iliac arteries were symmetrically modelled with an angle of 45◦ . The best fit circles and the centrelines were exported into Pro/Engineer Wildfire (v3, PTC Inc., MA, USA) as shown in Fig. 1(C). These circles were positioned concentrically on the straight centreline, thus creating a symmetrical model. Three surfaces were generated: the lumen, the ILT and the outer wall, identified in Fig. 1(D). The surfaces representing the lumen and thrombus were obtained by sweeping the generated circles. The third outer wall surface was generated by assuming a constant thickness of 2 mm, which was based on the literature values [8]. The full idealised model is presented in Fig. 1(D). CAD/CAM software, Pro/Manufacture (v3, PTC Inc., MA, USA), was used to generate the G codes necessary to manufacture three sets of modular moulds. The moulds were machined on a 3 axis CNC milling machine (Bridgeport Interact with a Fanuc controller). These modular moulds were employed to fabricate models with and without ILT, following the process shown in Fig. 2. The lost wax process and the injection method were used to create the physical
An experimental-flow-rig, shown in Fig. 3, was constructed to replicate in vivo conditions through the infrarenal AAA. Algorithms were designed and developed to control, monitor and record motion, data acquisition and vision systems by using LabVIEW (v8.6, National Instruments) software. A physiological flow waveform (Fig. 4) derived from the literature [27] was applied at the inlet. This waveform was replicated with a piston pump actuated by a computer programmable linear motor (LMA264, Aerotech, UK). Water–glycerine mixture was used as a blood analogue. The inlet waveform was monitored with an inline ultrasonic flowsensor (25PXN, Transonic Systems Inc., US). There was a 6.6% maximum difference in flowrate when the programmed input waveform was compared with the measured waveform as shown in Fig. 4. The internal pressure was monitored by a 3 French (Fr) pressure catheter (FTS-3011B, Scisense Inc.). The mean pressure (Pmean ) Table 1 Silicones used for each model and their resulted elastic modulus. Model
Components
Silicone
M1 M2
Wall Wall ILT-A Wall ILT-B Wall ILT-C
Elastosil 4641 Elastosil 4641 Elastosil 4600 Elastosil 4641 Elastosil 4600 Elastosil 4641 Elastosil 4600
M3 M4
Silicone fluid 5% 5% 70% 5% 50% 5% 25%
E [MPa]a 1.2 1.2 0.05 1.2 0.1 1.2 0.2
a Tensile testing done in-house conforms to the BS ISO 37:2005 standard (type 2, dog-bone) using an Instron tensile tester (model 5544) equipped with a 10N load cell and a standard video extensometer (SVE).
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Fig. 2. The manufacturing process of models with and without ILT. The same three sets of modular moulds were used for both models. (A) Mould set 1 creates the inner wax core by wax casting; (B) mould set 2 creates ILT by injecting ILT silicone for model with ILT or increases the size of the inner wax core by wax casting for model without ILT; (C) mould set 3 creates the outer wall by injecting Wall silicone for both models; (D) models with and without ILT are obtained after removing the wax core.
as defined by Eq. (1) [28] was kept at 100 mm Hg for all four models: Pmean =
Psys − 2Pdias 3
(1)
where Psys and Pdias are the systolic and diastolic pressure, respectively. The wall displacements were recorded by a vision system with a 4 megapixel resolution CCD camera (4M30, Dalsa Corporation; 15 frames per second) and an image pixel size of 0.049 mm. An automated straight edge detection tool tracked the required points during the pulse cycle. Local pressures and diameters were measured at four positions: the proximal aorta (position 1), the
proximal neck (position 2), the maximum aneurysm diameter (position 3) and the maximum thrombus area (position 4) as shown in Fig. 5. The cross sectional area of the thrombus in the physical model was 1400 mm2 (thickness 8.75 mm) at position 3 and varied along the aneurysm reaching maximum area of 1840 mm2 (thickness 14 mm) at position 4 as shown in Fig. 5. Acquisition and synchronisation of the flow, pressure and diameter data from the dynamic testing were achieved using a PXI system (PXI 1033, National Instruments, UK) with integrated high speed data acquisition and image acquisition modules. Representative flow, pressure and diameter temporal waveforms were obtained by fitting the data acquired over five pulse cycles with a Fourier series in trigonometric form.
Fig. 3. Experimental-flow-rig comprised of motion system that mimics blood flow, data acquisition that monitors the fluid’s parameters and the vision system that monitors the model’s parameters.
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As pressure, area and diameter vary during the pulse cycle the systolic components in Eqs. (2) and (3) were substituted with an instantaneous value for pressure, area and diameter, resulting in: Compliance (t) =
A(t) − Adias A(t)(P(t) − Pdias )
Diametral strain (t) =
Fig. 4. Validation of physiologic flow profile; input flow was obtained from the literature [27].
Compliance is a measure of the distensibility of the artery, and is defined [6] as: Compliance =
Asys − Adias Asys (Psys − Pdias )
(2)
where Psys and Pdias are the systolic and diastolic pressure, respectively; Asys and Adias are the systolic and diastolic area, respectively. Diametral strain over the pulse cycle is the change in diameter over the original diameter (the diastolic diameter): Diametral strain =
Dsys − Ddias Ddias
(3)
where Dsys and Ddias are the systolic and diastolic diameter, respectively.
Fig. 5. (A) Model of AAA without ILT and (B) model of AAA with ILT. Positions of diameter measurements: position 1 is the proximal aorta, position 2 is the proximal neck, position 3 is the maximum diameter and position 4 is the maximum thrombus area. At each position the external diameter (˚) and the ILT thickness (TILT ) are presented (all units are in mm).
D(t) − Ddias Ddias
(4)
(5)
where P(t), A(t) and D(t) as a function of time (t) are the instantaneous pressure, area and diameter values. Raghavan and Vorp [29] proposed a constitutive model for the AAA wall material based on the assumption of homogenous, incompressible, isotropic and hyperelastic material. From experimental tensile results of AAA specimens the population constitutive material properties ˛ = 174,000 ± 15,000 Pa and ˇ = 1,881,000 ± 372,000 Pa were estimated [29]. The mean (˛ = 174,000 Pa and ˇ = 1,881,000 Pa), minimum (˛min = 159,000 Pa and ˇmin = 1,509,000 Pa) and maximum (˛max = 189,000 Pa and ˇmax = 2,253,000 Pa) constitutive material properties together with the constitutive model [29] were used to validate the silicone used as an arterial mimicking material employing the finite element analysis (FEA) in ANSYS Structural (v11, ANSYS Inc., UK). A physical straight tube of diameter 20 mm and 2 mm wall thickness was created using the same silicone as the aortic wall of the models. A static test with interior pressures ranging from 0 to 150 mm Hg and a dynamic test were conducted for the straight tube. Pressure and diameter curves were measured during tests using the same procedure as describes above. 3. Results A straight tube was tested under static and physiological conditions and the pressure–diameter response of the silicone tube is shown in Fig. 6. The pressure–diameter curve was within the range as determined by the FEA (Fig. 6) based on the mean, minimal and maximal AAA material properties obtained by Raghavan and Vorp [29]. The static test shows a higher difference between the experimental and FEA results with only less than 1% variation in diameter when compared with the mean constitutive parameters (FEA − Mean). This validates that silicone is a suitable arterial mimicking material, within physiological pressure ranges. Pressure and diameter curves recorded for models M1 and M4, at position 3, are shown in Fig. 7. The diameter variation over the pulse cycle (Fig. 7(A and B)) for both types of models followed a similar profile to that of the pressure waveform (Fig. 7(A and B)). Fig. 7(C and D) presents hysteresis of pressure–diameter curves with loading and unloading cycles. The diameter is smaller on the expansion than during retraction. Local pressures were obtained for all four models maintaining the same input flow waveform and a mean pressure of 100 mm Hg. The variation in elastic stiffness for each model changed the recorded pressure waveform. The stiffer the model, a greater pressure difference (P) between systole and diastole (ranging from 17 to 45 mm Hg) was obtained, as seen in Fig. 8(A). The model M2 with the lowest ILT stiffness had relatively little effect on compliance as can be seen from the blood pressure waveform which was very similar to the model without ILT (M1), with a maximum percentage difference of 1%. The mean pressure varied slightly along the axial length for the stiffer models, as seen in Fig. 8(B). Table 2 shows the compliance calculated with Eq. (2) for all four models at positions defined in Fig. 5. The dynamic compliance of models M2, M3 and M4 was 13.87, 7.53 and 3.93 × 10−4 /mm Hg at maximum diameter, respectively and 16.3,
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Fig. 6. Comparison of pressure–diameter curves for silicone tube with FEA results using aneurismal material properties [29] for (A) static and (B) dynamic tests.
Fig. 7. Typical pressure/diameter versus time graphs (A and B) and pressure versus diameter graphs (C and D) for models M1 and M4.
8.40 and 3.09 × 10−4 /mmHg at position of maximum thrombus area, respectively. Compliance and P–D loading curves for the four models at four different locations are presented in Fig. 9. The instantaneous compliance profiles were calculated using Eq. (4).
Fig. 9(A1) shows a similar compliance for all the models at the proximal aorta (position 1). At the proximal neck, Fig. 9(A2) shows a lower compliance for the model without thrombus (M1) when compared with the models with thrombus (M2–M4). In the regions of maximum diameter and maximum ILT area as shown in Fig. 9(A3 and A4), model M4 had the lowest compli-
Fig. 8. (A) Pressure profiles response for realistic waveform for all models tested at position 1. Mean pressure kept at 100 mm Hg. (B) Longitudinal distribution of the mean pressure.
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Fig. 9. Compliance (A1–A4) and P–D curves (B1–B4) comparison at different positions along the models.
ance while models M1 and M2 had the highest compliance. The high thrombus area, shown in Fig. 9(A4), resulted in the highest difference in compliance between model M1 and models M3 and M4. Fig. 9(B1–B4) shows the P–D curves for the loading cycle. Fig. 9(B1 and B2) shows that the slope of the P–D curves is quite similar for all models, within a pressure change of 17 mm Hg. With the inclusion of the thrombus the slope of the P–D curves is
reduced for M3 and M4 as shown in Fig. 9(B3 and B4). For model M2 which has low thrombus stiffness the slope of the P–D curves was similar to the model without thrombus, M1. Fig. 10(A) shows the Gaussian curvature of the outer aneurismal wall for model without thrombus and the inner lumen wall for the model with thrombus. Fig. 10(B) shows the longitudinal distribution of average diametral strain. The two main regions of curvature change identified by the Gaussian curvature corresponded to the
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Fig. 10. (A) The Gaussian curvature for the outer aneurysm wall and the inner lumen wall; (B) longitudinal distribution of average diametral strains for all four models. (C–F) Diameter strain during the pulse cycle at the four positions for all models.
position of the highest diametral strain. Fig. 10(C–F) shows the diametral strain during the pulse cycle at the four positions for all models. For model M1 the highest diametral strain corresponded to the second region of outer wall curvature change (Fig. 10(C)). Model M2 also had the highest diametral strain at position 4 (Fig. 10(D)) which corresponded to the second region of outer and inner lumen wall curvature changes. For models M3 and M4 the highest diame-
tral strain corresponded with the first region of curvature change at the proximal neck of the models (Fig. 10(E and F)). 4. Discussion A flow rig was developed in this study to test the influence of ILT and its properties on an idealised aorto-iliac model
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Table 2 Compliance results for each model at different positions. M1
M2 −4
Compliance [10 Position 1 (proximal aorta) Position 2 (proximal neck) Position 3 (max diameter) Position 4 (max ILT area)
7.08 10.81 11.89 16.38
M3
M4
/mm Hg]
8.40 13.80 13.87 16.3
6.85 15.05 7.53 8.40
6.20 14.08 3.93 3.09
derived from a CT scan. Silicone was chosen as a modelling material as it offers similar mechanical properties, compliance and pressure–diameter response to that of an artery as shown in Fig. 6. The silicone AAA models have a compliant response to pressure and the pressure–diameter curves present hysteresis during the loading and unloading cycles. The diameter is smaller on the expansion than during retraction which is associated with hyperelastic materials, such as silicone, and also mimics aortic behaviour [30]. The dynamic compliance measured for the models with ILT varied from 3.09 to 16.30 × 10−4 /mm Hg. These results agree with the in vivo findings of Vorp et al. [6] in which compliance values varying from 1.8 to 9.4 × 10−4 /mm Hg were found, with an average of 4 × 10−4 /mm Hg. Also, Vorp et al. [6] observed that thrombus has a greater compliance than the abdominal aortic wall suggesting that ILT acts as a mechanical buffer or cushion. The role of intraluminal thrombus in AAA rupture and stress distribution is still not completely understood. Some in vivo studies have shown that the thrombus thickness was higher in patients with non-ruptured aneurysms as compared with ruptured aneurysms [31,32]. Pillari et al. [32] observed that the thrombus is generally thicker on the anterior wall of AAAs as compared with the rupture site at the posterior wall. Thubrikar et al. [33] showed, using in vivo and in vitro methods that the mural thrombus appears to protect the aneurysm by diminishing and homogenising strains in the aneurysm wall without restricting pressure transmission. Computational studies [7–12] of either simplified idealised models [10,12] or patient specific geometries of AAA with ILT [7–9,11] showed that, regardless of its mechanical properties, thrombus reduces and redistributes the stresses in the aneurismal wall markedly. The magnitude of the protective role of ILT depends on several factors like thrombus size, aneurysm size, thrombus properties, aneurysms stiffness and inhomogenities within thrombus [10,12]. Our results for the stiffer thrombi agree with these findings that thrombus may have a protective influence. Figs. 9 and 10 show that the diametral strains, compliance and P–D curves were influenced by the presence, elastic stiffness and thickness of the ILT. With the inclusion of ILT at maximum diameter (position 3) the lumen area of model M3 and M4 was reduced by 47% resulting in a 36–67% reduction in compliance compared to the model without ILT. While at maximum thrombus area (position 4) the lumen area was further reduced by 77% resulting in a 48 to 81% reduction in compliance. With an increase in elastic stiffness from 0.1 MPa for model M3 to 0.2 MPa for model M4, a reduction in compliance of 47 and 63% was observed for positions 3 and 4, respectively. The model with the lowest ILT stiffness (M2) had the opposite effect on compliance. There was no reduction in compliance in the distal region of the aneurismal sac and a 16% increase in compliance at position 3 for model M2 when compared with model M1. Fig. 10(A) shows three regions of curvature change. This occurred at the proximal neck, distal to the maximum diameter and at the aortic bifurcation. The region of curvature change situated distal to the maximum diameter induced maximum diametral strains in this region. With the inclusion of the ILT (M2–M4), it had a cushion effect in the regions of curvature change, thus reducing the diametral strains. Meanwhile, the ILT shifted the position of the
maximum diametral strain to the proximal neck (position 2) for the stiffer ILT models (M3 and M4) as shown in Fig. 10(B). When position 2 is compared to position 1, there was an increase in diametral strain of 40, 55 and 57% and an increase in compliance of 39, 55 and 56% for the models M2, M3 and M4, respectively. The lower stiffness of the ILT for M2 caused a prolonged increase in diametral strain for the region between proximal neck (position 2) and maximum ILT area (position 4). This was also due to the Gaussian curvature effect of the inner lumen wall and reduced ILT thickness along the proximal region. Beyond position 4, all models with ILT had a cushioning effect when compared with M1 (model without ILT). The ILT inclusion caused a material mismatch between the ILT and arterial wall to initiate at the proximal neck. This material mismatch and the region of curvature change found at the proximal neck may create another potential rupture site. This agrees with other studies that reported a rupture or the potential of rupture at proximal necks [15,34–37]. This work can be used for numerical validation purposes. While state of the art FE simulations can provide sophisticated modelling assumptions, its accuracy in terms of simulation parameters can influence numerical results quiet significantly. Our compliant models replicate an AAA geometry including iliac arteries and ILT with varying properties and complex distribution, which will help in validating the wall deformations and pressures for numerical studies. Our model had a symmetrical ILT distribution which contributed to the cushioning effect and shifting of the diametral strains. Some AAAs have a non-symmetrical ILT distribution. This would result in a material mismatch to occur at thrombus sites circumferentially along the model. The distribution of non-symmetrical ILT may explain the reported instances of the non-protective role of ILT on aneurysms. Pressure measurements during surgical exposure through fibrin thrombus failed to show a reduction in the pressure transmission on the wall. This indicates that ILT has no effect on aneurismal rupture [33,38,39]. Vorp et al. [40] suggested that ILT may act as a significant barrier to O2 transport from the lumen to the aortic wall, suggesting that hypoxia is present in regions of AAA wall adjacent to thick ILT [40], which may result in localised wall weakening and subsequent rupture. Our study did not take into consideration the effects of wall weakening or thinning in the region of ILT, which is a limitation in this work. A non-homogenous wall thickness in the presence of ILT has been shown to increase the wall stress and could predict aneurismal wall rupture better than a homogenous wall thickness [41,42]. The role of thrombus distribution and maturity on AAA mechanics was investigated in this work. However, the structure of in vivo intraluminal thrombus is inhomogeneous, fibrous and poro-elastic in nature [40,43] which may transmit the arterial pressure directly to the AAA wall. Silicone elastomers do not mimic the porosity or thrombus composition and this is a limitation of the present study. ILT properties vary somewhat in the radial direction. Wang et al. [4] and Gasser et al. [26] found that the ILT gets stiffer from luminal to abluminal layer. The inhomogeneous nature of the thrombus could indicate the transmission of stresses to the aneurismal wall in regions of thrombus failure resulting in the initiation of aneurismal wall failure [43–45]. We do not anticipate that these considerations greatly influence aneurysm mechanics as other in vitro studies that used human [23] and animal [46] thrombi to realistically represent its structure have suggested a protective effect of ILT, similar to our results. As an initial study the out of plane curvature was eliminated to assess the influence of thrombus. Future work will involve manufacturing a realistic model of AAA and studying the effects of other thrombus properties and non-symmetrical thrombus distribution in the aneurysm sac. An ultrasound imaging system will be added to the flow rig in order to monitor the internal wall displacements
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and the compressibility characteristic of the silicone material will also be assessed. This in vitro testing will also provide validation for other numerical studies of AAA. In conclusion, the results of this study showed a reduction of diametral strain, compliance and slope of P–D curves in models with mature ILT compared with the model without ILT in the region of maximum diameter, showing a beneficial impact of the thrombus on the mechanics of aneurysm rupture. The shifting of the maximum diametral strain and compliance to the proximal neck was pronounced by an increase in ILT stiffness.
Acknowledgement This project was funded by Enterprise Ireland under the Applied Research Enhancement (ARE) Programme (project code RE/05/009). Conflict of interest statement No conflict of interest.
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[24] Doyle BJ, Morris LG, Callanan A, Kelly P, Vorp DA, McGloughlin TM. 3D reconstruction and manufacture of real abdominal aortic aneurysms: from CT 480 scan to silicone model. Journal of Biomechanical Engineering 2008;130(Jun (3)):034501. [25] Timaran CH, Ohki T, Veith FJ, Lipsitz EC, Gargiulo NJ, Rhee SJ, et al. Influence of type II endoleak volume on aneurysm wall pressure and distribution in an experimental model. Journal of Vascular Surgery 2005;41(4):657–63. [26] Gasser TC, Gorgulu G, Folkesson M, Swedenborg J. Failure properties of intraluminal thrombus in abdominal aortic aneurysm under static and pulsating mechanical loads. Journal of Vascular Surgery 2008;48(July (1)): 179–88. [27] Li Z, Kleinstreuer C. Blood flow and structure interactions in a stented abdominal aortic aneurysm model. Medical Engineering & Physics 2005;27(5): 369–82. [28] Klabunde RE. Cardiovascular Physiology Concepts. Lippincott: Williams & Wilkins; 2005. [29] Raghavan ML, Vorp DA. Toward a biomechanical tool to evaluate rupture potential of abdominal aortic aneurysm: identification of a finite strain constitutive model and evaluation of its applicability. Journal of Biomechanics 2000;33(April (4)):475–82. [30] Sonesson B, Lanne T, Vernersson E, Hansen F. Sex difference in the mechanical properties of the abdominal aorta in human beings. Journal of Vascular Surgery 1994;20(December (6)):959–69. [31] Kushihashi T, Munechika H, Matsui S, Moritani T, Horichi Y, Hishida T. CT of abdominal aortic aneurysms—aneurysmal size and thickness of intraaneurysmal thrombus as risk factors of rupture. Nihon Igaku Hoshasen Gakkai zasshi 1991;51(March (3)):219–27. [32] Pillari G, Chang JB, Zito J, Cohen JR, Gersten K, Rizzo A, et al. Computed tomography of abdominal aortic aneurysm. An in vivo pathological report with a note on dynamic predictors Archives of Surgery 1988;123(6):727–32. [33] Thubrikar MJ, Robicsek F, Labrosse M, Chervenkoff V, Fowler BL. Effect of thrombus on abdominal aortic aneurysm wall dilation and stress. Journal of Cardiovascular Surgery 2003;44(February (1)):67–77. [34] Fillinger MF, Marra SP, Raghavan ML, Kennedy FE. Prediction of rupture risk in abdominal aortic aneurysm during observation: wall stress versus diameter. Journal of Vascular Surgery 2003;37(4):724–32. [35] Morris L, O’Donnell P, Delassus P, McGloughlin T. Experimental assessment of stress patterns in abdominal aortic aneurysms using the photoelastic method. Strain 2004;40(4):165–72. [36] Raghavan ML, Vorp DA, Federle MP, Makaroun MS, Webster MW. Wall stress distribution on three-dimensionally reconstructed models of human abdominal aortic aneurysm. Journal of Vascular Surgery 2000;31(4):760–9. [37] Sacks MS, Vorp DA, Raghavan ML, Federle MP, Webster MW. In vivo threedimensional surface geometry of abdominal aortic aneurysms. Annals of Biomedical Engineering 1999;27(4):469–79. [38] Schurink GWH, Baalen JMv, Visser MJT, Bockel JHv. Thrombus within an aortic aneurysm does not reduce pressure on the aneurysmal wall. Journal of Vascular Surgery: Official Publication, The Society for Vascular Surgery [and] International Society for Cardiovascular Surgery, North American Chapter 2000;31(3):501–6. [39] Hans SS, Jareunpoon O, Huang R, Hans B, Bove P, Zelenock GB. Relationship of residual intraluminal to intrathrombotic pressure in a closed aneurysmal sac. Journal of Vascular Surgery: Official Publication, The Society for Vascular Surgery [and] International Society for Cardiovascular Surgery, North American Chapter 2003;37(5):949–53. [40] Vorp DA, Lee PC, Wang DH, Makaroun MS, Nemoto EM, Ogawa S, et al. Association of intraluminal thrombus in abdominal aortic aneurysm with local hypoxia and wall weakening. Journal of Vascular Surgery 2001;34(2):291–9. [41] Gasser TC, Auer M, Labruto F, Swedenborg J, Roy J. Biomechanical Rupture Risk Assessment of Abdominal Aortic Aneurysms: Model Complexity versus Predictability of Finite Element Simulations. European Journal of Vascular and Endovascular Surgery 2010;40(2):176–85. [42] Raghavan ML, Kratzberg J, Castro de Tolosa EM, Hanaoka MM, Walker P, da Silva ES. Regional distribution of wall thickness and failure properties of human abdominal aortic aneurysm. Journal of Biomechanics 2006;39(16): 3010–6. [43] Gasser TC, Martufi G, Auer M, Folkesson M, Swedenborg J. Micromechanical characterization of intra-luminal thrombus tissue from abdominal aortic aneurysms. Annals of Biomedical Engineering 2009 November 17. [44] Kazi M, Thyberg J, Religa P, Roy J, Eriksson P, Hedin U, et al. Influence of intraluminal thrombus on structural and cellular composition of abdominal aortic aneurysm wall. Journal of Vascular Surgery 2003;38(6): 1283–92.
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Medical Engineering & Physics journal homepage: www.elsevier.com/locate/medengphy
In vitro evaluation of the effects of intraluminal thrombus on abdominal aortic aneurysm wall dynamics Florentina Ene ∗ , Carine Gachon, Patrick Delassus, Ronan Carroll, Florian Stefanov, Padraig O’Flynn, Liam Morris Department of Mechanical Engineering, Galway Medical Technologies Centre (GMedTech), Galway Mayo Institute of Technology (GMIT), Dublin Road, Galway, Ireland
a r t i c l e
i n f o
Article history: Received 20 January 2010 Received in revised form 12 March 2011 Accepted 14 March 2011 Keywords: Abdominal aortic aneurysm Intraluminal thrombus Silicone models Compliance Diametral strain
a b s t r a c t The optimum time to treat abdominal aortic aneurysms (AAAs) still remains an uncertain issue. The decision to intervene does not take in account the effects that wall curvature, intraluminal thrombus (ILT) properties and thickness have on rupture. The role of ILT in aneurysm dynamics and rupture has been controversial. In vitro testing of four silicone AAA models incorporating the ILT and aortic bifurcation was studied under physiological conditions. Pressures (P) and diameters (D) were analysed for models with and without ILT at different locations. The diametral strain, compliance and P/D curves were influenced by the presence, elastic stiffness and thickness of the ILT. In this case, the inclusion of ILT reduced the lumen area by 77% that resulted in a 0.5–81% reduction in compliance depending on ILT properties. With an increase in ILT stiffness from 0.05 to 0.2 MPa, the compliance was reduced by 81%. In the region of maximum diameter, there was a reduction of diametral strain and compliance except for the softer ILT which was more compliant throughout the proximal region. The shifting of the maximum diametral strain and compliance to the proximal neck was pronounced by an increase in ILT stiffness, thus creating a possible rupture site. © 2011 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction Abdominal aortic aneurysm (AAA) is one of the leading causes of death in the western world, being the 14th cause of death in the U.S. for men over the age of 55 [1,2]. An aneurysm is an area of localised and permanent dilatation of a blood vessel. Aortic aneurysms can develop anywhere along the length of the aorta, but the majority are located below the renal arteries and are called abdominal aortic aneurysms (AAAs). These AAAs can also extend into one or both of the iliac arteries. The morphology of the bulging in an AAA is patient specific and most commonly takes a fusiform shape. Aneurysms have a high risk of rupture if not detected and treated in time. The optimum time to treat AAAs still remains an uncertain issue. The majority of aneurysms (75%) contain intraluminal thrombus (ILT), however the extent of its presence varies [3]. The thrombus is a fibrin structure incorporated with blood cells, platelets, blood proteins and cellular debris [3]. Mechanical testing has proven that the ILT is a nonlinear elastic isotropic and inhomogeneous material [4]. Its exact role in AAA rupture has been controversial. Examination of ruptured AAAs has shown the presence of ILT at the location of failure [5]. It has been proposed that the presence
∗ Corresponding author. Tel.: +353 862050346. E-mail address: [email protected] (F. Ene).
of ILT induces hypoxia thus weakening the aneurismal wall and contributing to AAA rupture [40]. Conversely, other studies have proven that the ILT acts like a mechanical cushion with smaller stress in the underlying AAA wall [6], therefore possibly protecting it from rupture [7]. Computational studies of wall stress distribution have shown that, regardless of its mechanical properties, thrombus markedly reduces and redistributes the stresses in the aneurismal wall [7–12]. In vitro testing is an alternative to computational work for studying aneurysm mechanics and blood flow dynamics. Important developments have been made over the years to recreate the physiological conditions [13–17]. Compliant vascular replicas have been produced by researchers using arterial mimicking materials such as latex, silicone, polymers, gelatine and other materials [15,16,18–22]. Aneurismal models have included fusiform [15,23], idealised [16,24] and realistic with the iliac arteries [16,18]. These models however neglected to account for ILT. In a small number of studies the intraluminal thrombus has been approximated using gelatine solution [18] and dough [25], with no reported mechanical properties, inserted in dipped latex models. Other studies replicated ILT using customised polymer mixtures [21] and silicone [47]. To date, no experimental study has been done on the effect that ILT has on AAA wall dynamics, in terms of compliance and pressure/diameter variations. The objective of the current study was
1350-4533/$ – see front matter © 2011 IPEM. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.medengphy.2011.03.005
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Fig. 1. The planar idealised model of AAA. (A) The lumen of AAA with iliac arteries as reconstructed from CT scan; (B) an individual CT slice at level of aneurysm presenting the lumen region and the thrombus region from which the thrombus geometry was obtained; (C) the best fit circles of the lumen with a straight centreline; (D) the final idealised abdominal aorta with aneurysm, ILT and iliac arteries.
to analyse experimentally the effects of pulsating pressure on an idealised aneurismal aorto-iliac model with and without intraluminal thrombus. Influencing factors contributing to the complex AAA mechanics such as thrombus stiffness, thrombus distribution throughout the aneurysm sac and the aneurismal wall curvature were evaluated. The specific aims were to evaluate compliance, diametral strain and pressure/diameter variations under physiological conditions and identify the effects of ILT on wall dynamics. The results of this study can be applied to validate numerical studies in complex AAA geometries. 2. Methods
models. The wax (Texwax FG890, Texaco, Germany) created the lumen cavity. The silicones used for the ILT and arterial wall are listed in Table 1. These silicones (Wacker Chemie AG, Germany) were translucent two-component (A and B) silicone elastomers that was mixed in 10:1 ratio. Silicone fluid (Dow Corning, UK) was added in different percentages to lower the elastic modulus. For this study, four AAA models were created, one model without thrombus (M1) and three models with thrombus (M2–M4) as presented in Table 1. Three different thrombus properties were modelled, with elastic modulus of 0.05 MPa (M2), 0.1 MPa (M3) and 0.2 MPa (M4). These elastic values showing different ILT maturity are within the range (0.04–0.27 MPa) documented in the literature [4,8,26].
2.1. Physical model generation
2.2. Dynamic physiological testing of AAA models
A computed tomography (CT) scan of a male subject, suffering from an abdominal aortic aneurysm was obtained from Materialise Medical Datasets (Leuven, Belgium). The scans had a slice thickness of 5 mm with a pixel size of 0.8242 mm. A three-dimensional (3D) aortic model was created using Mimics (v11, Materialise, Leuven, Belgium) through the process of segmentation. This model exhibited non-planar curvature due to its course through the abdomen as shown in Fig. 1(A). A planar idealised model was generated by eliminating the effects of the curvature. At each scan slice, the lumen and the thrombus areas were generated (Fig. 1(B)) and a best fit circle was obtained for each area. The realistic vessel centrelines were converted to straight lines of equal length. The iliac arteries were symmetrically modelled with an angle of 45◦ . The best fit circles and the centrelines were exported into Pro/Engineer Wildfire (v3, PTC Inc., MA, USA) as shown in Fig. 1(C). These circles were positioned concentrically on the straight centreline, thus creating a symmetrical model. Three surfaces were generated: the lumen, the ILT and the outer wall, identified in Fig. 1(D). The surfaces representing the lumen and thrombus were obtained by sweeping the generated circles. The third outer wall surface was generated by assuming a constant thickness of 2 mm, which was based on the literature values [8]. The full idealised model is presented in Fig. 1(D). CAD/CAM software, Pro/Manufacture (v3, PTC Inc., MA, USA), was used to generate the G codes necessary to manufacture three sets of modular moulds. The moulds were machined on a 3 axis CNC milling machine (Bridgeport Interact with a Fanuc controller). These modular moulds were employed to fabricate models with and without ILT, following the process shown in Fig. 2. The lost wax process and the injection method were used to create the physical
An experimental-flow-rig, shown in Fig. 3, was constructed to replicate in vivo conditions through the infrarenal AAA. Algorithms were designed and developed to control, monitor and record motion, data acquisition and vision systems by using LabVIEW (v8.6, National Instruments) software. A physiological flow waveform (Fig. 4) derived from the literature [27] was applied at the inlet. This waveform was replicated with a piston pump actuated by a computer programmable linear motor (LMA264, Aerotech, UK). Water–glycerine mixture was used as a blood analogue. The inlet waveform was monitored with an inline ultrasonic flowsensor (25PXN, Transonic Systems Inc., US). There was a 6.6% maximum difference in flowrate when the programmed input waveform was compared with the measured waveform as shown in Fig. 4. The internal pressure was monitored by a 3 French (Fr) pressure catheter (FTS-3011B, Scisense Inc.). The mean pressure (Pmean ) Table 1 Silicones used for each model and their resulted elastic modulus. Model
Components
Silicone
M1 M2
Wall Wall ILT-A Wall ILT-B Wall ILT-C
Elastosil 4641 Elastosil 4641 Elastosil 4600 Elastosil 4641 Elastosil 4600 Elastosil 4641 Elastosil 4600
M3 M4
Silicone fluid 5% 5% 70% 5% 50% 5% 25%
E [MPa]a 1.2 1.2 0.05 1.2 0.1 1.2 0.2
a Tensile testing done in-house conforms to the BS ISO 37:2005 standard (type 2, dog-bone) using an Instron tensile tester (model 5544) equipped with a 10N load cell and a standard video extensometer (SVE).
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Fig. 2. The manufacturing process of models with and without ILT. The same three sets of modular moulds were used for both models. (A) Mould set 1 creates the inner wax core by wax casting; (B) mould set 2 creates ILT by injecting ILT silicone for model with ILT or increases the size of the inner wax core by wax casting for model without ILT; (C) mould set 3 creates the outer wall by injecting Wall silicone for both models; (D) models with and without ILT are obtained after removing the wax core.
as defined by Eq. (1) [28] was kept at 100 mm Hg for all four models: Pmean =
Psys − 2Pdias 3
(1)
where Psys and Pdias are the systolic and diastolic pressure, respectively. The wall displacements were recorded by a vision system with a 4 megapixel resolution CCD camera (4M30, Dalsa Corporation; 15 frames per second) and an image pixel size of 0.049 mm. An automated straight edge detection tool tracked the required points during the pulse cycle. Local pressures and diameters were measured at four positions: the proximal aorta (position 1), the
proximal neck (position 2), the maximum aneurysm diameter (position 3) and the maximum thrombus area (position 4) as shown in Fig. 5. The cross sectional area of the thrombus in the physical model was 1400 mm2 (thickness 8.75 mm) at position 3 and varied along the aneurysm reaching maximum area of 1840 mm2 (thickness 14 mm) at position 4 as shown in Fig. 5. Acquisition and synchronisation of the flow, pressure and diameter data from the dynamic testing were achieved using a PXI system (PXI 1033, National Instruments, UK) with integrated high speed data acquisition and image acquisition modules. Representative flow, pressure and diameter temporal waveforms were obtained by fitting the data acquired over five pulse cycles with a Fourier series in trigonometric form.
Fig. 3. Experimental-flow-rig comprised of motion system that mimics blood flow, data acquisition that monitors the fluid’s parameters and the vision system that monitors the model’s parameters.
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As pressure, area and diameter vary during the pulse cycle the systolic components in Eqs. (2) and (3) were substituted with an instantaneous value for pressure, area and diameter, resulting in: Compliance (t) =
A(t) − Adias A(t)(P(t) − Pdias )
Diametral strain (t) =
Fig. 4. Validation of physiologic flow profile; input flow was obtained from the literature [27].
Compliance is a measure of the distensibility of the artery, and is defined [6] as: Compliance =
Asys − Adias Asys (Psys − Pdias )
(2)
where Psys and Pdias are the systolic and diastolic pressure, respectively; Asys and Adias are the systolic and diastolic area, respectively. Diametral strain over the pulse cycle is the change in diameter over the original diameter (the diastolic diameter): Diametral strain =
Dsys − Ddias Ddias
(3)
where Dsys and Ddias are the systolic and diastolic diameter, respectively.
Fig. 5. (A) Model of AAA without ILT and (B) model of AAA with ILT. Positions of diameter measurements: position 1 is the proximal aorta, position 2 is the proximal neck, position 3 is the maximum diameter and position 4 is the maximum thrombus area. At each position the external diameter (˚) and the ILT thickness (TILT ) are presented (all units are in mm).
D(t) − Ddias Ddias
(4)
(5)
where P(t), A(t) and D(t) as a function of time (t) are the instantaneous pressure, area and diameter values. Raghavan and Vorp [29] proposed a constitutive model for the AAA wall material based on the assumption of homogenous, incompressible, isotropic and hyperelastic material. From experimental tensile results of AAA specimens the population constitutive material properties ˛ = 174,000 ± 15,000 Pa and ˇ = 1,881,000 ± 372,000 Pa were estimated [29]. The mean (˛ = 174,000 Pa and ˇ = 1,881,000 Pa), minimum (˛min = 159,000 Pa and ˇmin = 1,509,000 Pa) and maximum (˛max = 189,000 Pa and ˇmax = 2,253,000 Pa) constitutive material properties together with the constitutive model [29] were used to validate the silicone used as an arterial mimicking material employing the finite element analysis (FEA) in ANSYS Structural (v11, ANSYS Inc., UK). A physical straight tube of diameter 20 mm and 2 mm wall thickness was created using the same silicone as the aortic wall of the models. A static test with interior pressures ranging from 0 to 150 mm Hg and a dynamic test were conducted for the straight tube. Pressure and diameter curves were measured during tests using the same procedure as describes above. 3. Results A straight tube was tested under static and physiological conditions and the pressure–diameter response of the silicone tube is shown in Fig. 6. The pressure–diameter curve was within the range as determined by the FEA (Fig. 6) based on the mean, minimal and maximal AAA material properties obtained by Raghavan and Vorp [29]. The static test shows a higher difference between the experimental and FEA results with only less than 1% variation in diameter when compared with the mean constitutive parameters (FEA − Mean). This validates that silicone is a suitable arterial mimicking material, within physiological pressure ranges. Pressure and diameter curves recorded for models M1 and M4, at position 3, are shown in Fig. 7. The diameter variation over the pulse cycle (Fig. 7(A and B)) for both types of models followed a similar profile to that of the pressure waveform (Fig. 7(A and B)). Fig. 7(C and D) presents hysteresis of pressure–diameter curves with loading and unloading cycles. The diameter is smaller on the expansion than during retraction. Local pressures were obtained for all four models maintaining the same input flow waveform and a mean pressure of 100 mm Hg. The variation in elastic stiffness for each model changed the recorded pressure waveform. The stiffer the model, a greater pressure difference (P) between systole and diastole (ranging from 17 to 45 mm Hg) was obtained, as seen in Fig. 8(A). The model M2 with the lowest ILT stiffness had relatively little effect on compliance as can be seen from the blood pressure waveform which was very similar to the model without ILT (M1), with a maximum percentage difference of 1%. The mean pressure varied slightly along the axial length for the stiffer models, as seen in Fig. 8(B). Table 2 shows the compliance calculated with Eq. (2) for all four models at positions defined in Fig. 5. The dynamic compliance of models M2, M3 and M4 was 13.87, 7.53 and 3.93 × 10−4 /mm Hg at maximum diameter, respectively and 16.3,
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Fig. 6. Comparison of pressure–diameter curves for silicone tube with FEA results using aneurismal material properties [29] for (A) static and (B) dynamic tests.
Fig. 7. Typical pressure/diameter versus time graphs (A and B) and pressure versus diameter graphs (C and D) for models M1 and M4.
8.40 and 3.09 × 10−4 /mmHg at position of maximum thrombus area, respectively. Compliance and P–D loading curves for the four models at four different locations are presented in Fig. 9. The instantaneous compliance profiles were calculated using Eq. (4).
Fig. 9(A1) shows a similar compliance for all the models at the proximal aorta (position 1). At the proximal neck, Fig. 9(A2) shows a lower compliance for the model without thrombus (M1) when compared with the models with thrombus (M2–M4). In the regions of maximum diameter and maximum ILT area as shown in Fig. 9(A3 and A4), model M4 had the lowest compli-
Fig. 8. (A) Pressure profiles response for realistic waveform for all models tested at position 1. Mean pressure kept at 100 mm Hg. (B) Longitudinal distribution of the mean pressure.
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Fig. 9. Compliance (A1–A4) and P–D curves (B1–B4) comparison at different positions along the models.
ance while models M1 and M2 had the highest compliance. The high thrombus area, shown in Fig. 9(A4), resulted in the highest difference in compliance between model M1 and models M3 and M4. Fig. 9(B1–B4) shows the P–D curves for the loading cycle. Fig. 9(B1 and B2) shows that the slope of the P–D curves is quite similar for all models, within a pressure change of 17 mm Hg. With the inclusion of the thrombus the slope of the P–D curves is
reduced for M3 and M4 as shown in Fig. 9(B3 and B4). For model M2 which has low thrombus stiffness the slope of the P–D curves was similar to the model without thrombus, M1. Fig. 10(A) shows the Gaussian curvature of the outer aneurismal wall for model without thrombus and the inner lumen wall for the model with thrombus. Fig. 10(B) shows the longitudinal distribution of average diametral strain. The two main regions of curvature change identified by the Gaussian curvature corresponded to the
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Fig. 10. (A) The Gaussian curvature for the outer aneurysm wall and the inner lumen wall; (B) longitudinal distribution of average diametral strains for all four models. (C–F) Diameter strain during the pulse cycle at the four positions for all models.
position of the highest diametral strain. Fig. 10(C–F) shows the diametral strain during the pulse cycle at the four positions for all models. For model M1 the highest diametral strain corresponded to the second region of outer wall curvature change (Fig. 10(C)). Model M2 also had the highest diametral strain at position 4 (Fig. 10(D)) which corresponded to the second region of outer and inner lumen wall curvature changes. For models M3 and M4 the highest diame-
tral strain corresponded with the first region of curvature change at the proximal neck of the models (Fig. 10(E and F)). 4. Discussion A flow rig was developed in this study to test the influence of ILT and its properties on an idealised aorto-iliac model
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Table 2 Compliance results for each model at different positions. M1
M2 −4
Compliance [10 Position 1 (proximal aorta) Position 2 (proximal neck) Position 3 (max diameter) Position 4 (max ILT area)
7.08 10.81 11.89 16.38
M3
M4
/mm Hg]
8.40 13.80 13.87 16.3
6.85 15.05 7.53 8.40
6.20 14.08 3.93 3.09
derived from a CT scan. Silicone was chosen as a modelling material as it offers similar mechanical properties, compliance and pressure–diameter response to that of an artery as shown in Fig. 6. The silicone AAA models have a compliant response to pressure and the pressure–diameter curves present hysteresis during the loading and unloading cycles. The diameter is smaller on the expansion than during retraction which is associated with hyperelastic materials, such as silicone, and also mimics aortic behaviour [30]. The dynamic compliance measured for the models with ILT varied from 3.09 to 16.30 × 10−4 /mm Hg. These results agree with the in vivo findings of Vorp et al. [6] in which compliance values varying from 1.8 to 9.4 × 10−4 /mm Hg were found, with an average of 4 × 10−4 /mm Hg. Also, Vorp et al. [6] observed that thrombus has a greater compliance than the abdominal aortic wall suggesting that ILT acts as a mechanical buffer or cushion. The role of intraluminal thrombus in AAA rupture and stress distribution is still not completely understood. Some in vivo studies have shown that the thrombus thickness was higher in patients with non-ruptured aneurysms as compared with ruptured aneurysms [31,32]. Pillari et al. [32] observed that the thrombus is generally thicker on the anterior wall of AAAs as compared with the rupture site at the posterior wall. Thubrikar et al. [33] showed, using in vivo and in vitro methods that the mural thrombus appears to protect the aneurysm by diminishing and homogenising strains in the aneurysm wall without restricting pressure transmission. Computational studies [7–12] of either simplified idealised models [10,12] or patient specific geometries of AAA with ILT [7–9,11] showed that, regardless of its mechanical properties, thrombus reduces and redistributes the stresses in the aneurismal wall markedly. The magnitude of the protective role of ILT depends on several factors like thrombus size, aneurysm size, thrombus properties, aneurysms stiffness and inhomogenities within thrombus [10,12]. Our results for the stiffer thrombi agree with these findings that thrombus may have a protective influence. Figs. 9 and 10 show that the diametral strains, compliance and P–D curves were influenced by the presence, elastic stiffness and thickness of the ILT. With the inclusion of ILT at maximum diameter (position 3) the lumen area of model M3 and M4 was reduced by 47% resulting in a 36–67% reduction in compliance compared to the model without ILT. While at maximum thrombus area (position 4) the lumen area was further reduced by 77% resulting in a 48 to 81% reduction in compliance. With an increase in elastic stiffness from 0.1 MPa for model M3 to 0.2 MPa for model M4, a reduction in compliance of 47 and 63% was observed for positions 3 and 4, respectively. The model with the lowest ILT stiffness (M2) had the opposite effect on compliance. There was no reduction in compliance in the distal region of the aneurismal sac and a 16% increase in compliance at position 3 for model M2 when compared with model M1. Fig. 10(A) shows three regions of curvature change. This occurred at the proximal neck, distal to the maximum diameter and at the aortic bifurcation. The region of curvature change situated distal to the maximum diameter induced maximum diametral strains in this region. With the inclusion of the ILT (M2–M4), it had a cushion effect in the regions of curvature change, thus reducing the diametral strains. Meanwhile, the ILT shifted the position of the
maximum diametral strain to the proximal neck (position 2) for the stiffer ILT models (M3 and M4) as shown in Fig. 10(B). When position 2 is compared to position 1, there was an increase in diametral strain of 40, 55 and 57% and an increase in compliance of 39, 55 and 56% for the models M2, M3 and M4, respectively. The lower stiffness of the ILT for M2 caused a prolonged increase in diametral strain for the region between proximal neck (position 2) and maximum ILT area (position 4). This was also due to the Gaussian curvature effect of the inner lumen wall and reduced ILT thickness along the proximal region. Beyond position 4, all models with ILT had a cushioning effect when compared with M1 (model without ILT). The ILT inclusion caused a material mismatch between the ILT and arterial wall to initiate at the proximal neck. This material mismatch and the region of curvature change found at the proximal neck may create another potential rupture site. This agrees with other studies that reported a rupture or the potential of rupture at proximal necks [15,34–37]. This work can be used for numerical validation purposes. While state of the art FE simulations can provide sophisticated modelling assumptions, its accuracy in terms of simulation parameters can influence numerical results quiet significantly. Our compliant models replicate an AAA geometry including iliac arteries and ILT with varying properties and complex distribution, which will help in validating the wall deformations and pressures for numerical studies. Our model had a symmetrical ILT distribution which contributed to the cushioning effect and shifting of the diametral strains. Some AAAs have a non-symmetrical ILT distribution. This would result in a material mismatch to occur at thrombus sites circumferentially along the model. The distribution of non-symmetrical ILT may explain the reported instances of the non-protective role of ILT on aneurysms. Pressure measurements during surgical exposure through fibrin thrombus failed to show a reduction in the pressure transmission on the wall. This indicates that ILT has no effect on aneurismal rupture [33,38,39]. Vorp et al. [40] suggested that ILT may act as a significant barrier to O2 transport from the lumen to the aortic wall, suggesting that hypoxia is present in regions of AAA wall adjacent to thick ILT [40], which may result in localised wall weakening and subsequent rupture. Our study did not take into consideration the effects of wall weakening or thinning in the region of ILT, which is a limitation in this work. A non-homogenous wall thickness in the presence of ILT has been shown to increase the wall stress and could predict aneurismal wall rupture better than a homogenous wall thickness [41,42]. The role of thrombus distribution and maturity on AAA mechanics was investigated in this work. However, the structure of in vivo intraluminal thrombus is inhomogeneous, fibrous and poro-elastic in nature [40,43] which may transmit the arterial pressure directly to the AAA wall. Silicone elastomers do not mimic the porosity or thrombus composition and this is a limitation of the present study. ILT properties vary somewhat in the radial direction. Wang et al. [4] and Gasser et al. [26] found that the ILT gets stiffer from luminal to abluminal layer. The inhomogeneous nature of the thrombus could indicate the transmission of stresses to the aneurismal wall in regions of thrombus failure resulting in the initiation of aneurismal wall failure [43–45]. We do not anticipate that these considerations greatly influence aneurysm mechanics as other in vitro studies that used human [23] and animal [46] thrombi to realistically represent its structure have suggested a protective effect of ILT, similar to our results. As an initial study the out of plane curvature was eliminated to assess the influence of thrombus. Future work will involve manufacturing a realistic model of AAA and studying the effects of other thrombus properties and non-symmetrical thrombus distribution in the aneurysm sac. An ultrasound imaging system will be added to the flow rig in order to monitor the internal wall displacements
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and the compressibility characteristic of the silicone material will also be assessed. This in vitro testing will also provide validation for other numerical studies of AAA. In conclusion, the results of this study showed a reduction of diametral strain, compliance and slope of P–D curves in models with mature ILT compared with the model without ILT in the region of maximum diameter, showing a beneficial impact of the thrombus on the mechanics of aneurysm rupture. The shifting of the maximum diametral strain and compliance to the proximal neck was pronounced by an increase in ILT stiffness.
Acknowledgement This project was funded by Enterprise Ireland under the Applied Research Enhancement (ARE) Programme (project code RE/05/009). Conflict of interest statement No conflict of interest.
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