Publisher’s version / Version de l'éditeur:
Journal of Materials Science (Letters), 40, 9-10, pp. 2675-2677, 2005-05-01
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How To Design a Structure Able to Mimic the Arterial Wall Mechanical
Behavior?
Jouan, M. R.; Bureau, M. N.; Ajji, A.; Huneault, M. A.; Mantovani, D.
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https://nrc-publications.canada.ca/eng/view/object/?id=7e84f565-05de-432b-b58c-beb763313d50 https://publications-cnrc.canada.ca/fra/voir/objet/?id=7e84f565-05de-432b-b58c-beb763313d50
J O U R N A L O F M A T E R I A L S S C I E N C E 4 0 (2 0 0 5 ) 2675 – 2677 L E T T E R S
How to design a structure able to mimic the arterial wall
mechanical behavior?
M . R . J O U A N
Department of Material Engineering, Laval University, Qu ´ebec, Canada; Unit of Bioengineering and Biotechnology, St Fran ¸cois d’Assise Hospital, Qu ´ebec, Canada
M . N . B U R E A U , A . A J J I , M . A . H U N E A U L T
Industrial Materials Institute, National Research Council, Boucherville, Canada
D. MANTOVANI
Department of Material Engineering, Laval University, Qu ´ebec, Canada; Unit of Bioengineering and Biotechnology, St Fran ¸cois d’Assise Hospital, Qu ´ebec, Canada
Cardiovascular diseases, particularly arterioscleroses, count among the most frequent and fatal ones in the industrialized countries [1]. Arteriosclerosis is caused by the formation of atheromateous plate constituted by proliferation and deposits of cholesterol in lipidic scratches involving thickening of the intima, which leads to a stenosis. Clinical complications are also re-lated to the possibility of calcification eventually lead-ing to a rupture, cholesterol embolus and ulceration of the thrombotic plate. From a clinical point of view, it is often required to replace the diseased and struc-turally damaged section of the artery and implant a vascular substitute, made from either synthetic (Teflon or Dacron prosthesis) or biological (autologous graft) materials. These substitutes must conform to several requirements such as hemocompatibility, mechanical reliability and iso-compliance, i.e., graft compliance as similar as possible to that of arteries to ensure their har-monious implantation. This compliance requirement is essential to obtain similar deformation under pressure in the prosthesis and adjacent artery. Failure to obtain acceptable compliance may result in improper anasto-mosis and mechanical mismatch, which eventually lead to the failure of the implant [2]. Several studies have fo-cused on the compliance of arteries to better understand their behavior [3–6].
Quite compelling insights on successful vascular grafts can also be obtained from a good understand-ing of the architecture of the arterial wall. In fact, the arterial wall is composed of three principal layers with very specific properties [7, 8]. The intima, the internal layer of the arterial wall (∼10% thickness of the wall), is composed of a unique layer of endothelial cells and serves as an internal artery liner reducing shearing of blood on the wall. It is not considered to contribute sig-nificantly to the elastic properties of the artery. The me-dia, the intermediate layer of the arterial wall (∼70% thickness of the wall), is the structural component of the artery. It is composed of elastin and collagen fibers circumferentially organized that define most the elas-tic properties of the artery. Smooth muscles are also present in muscular arteries, but their mechanical prop-erties can be considered as negligible. The adventitia,
the external layer of the arterial wall (∼20% thickness of the wall), is composed of elastin and collagen fibers and contributes to the elastic properties of the artery, although to a lesser extent than the media does. The principal role of this layer is to allow anchoring to the surrounding tissues.
Since the media accounts for most of the elastic prop-erties of the artery, it is considered to play a key role in regard to the elastic properties of the artery. The elastic moduli of the elastin and collagen fibers are reported to be approximately 3 and 1000 MPa, respectively [9]. The elastin fibers form a network of random macro-molecular entanglements that constitute the matrix of the media. The collagen fibers form a 3D network of aligned fibers within the elastin matrix in the media. Due to their large differences in modulus and structure, the resulting pressure vs. deformation curve for the me-dia, which constitutes the compliance curve, shows a characteristic concave shape, also called “J” shape [10], schematized in Fig. 1. As suggested in Fig. 1, this char-acteristic shape is caused by the sequential deformation of the elastin and collagen fibers upon loading. First, elastin fibers undergo re-orientation into the principal directions by rotation and translation due to their low modulus and random orientation. This results in the first part of the compliance curve characterized by a slope E1characteristic of the elastin property with a limited
contribution from the collagen fibers. Upon deforma-tion, the 3D network of collagen fibers is progressively oriented, which increases the stiffness of the material in the deformation direction and results in an increas-ing slope up to the E2 value of the collagen fibers. It
thus appears that, when the pressure increases in the artery, the wall is submitted to tension, which leads to artery radius increases up to the structural limit im-posed by the stronger and more rigid collagen fiber network.
A striking similarity between this behavior of the ar-teries and the behavior of orthogonal fabrics made from rovings of structural continuous fiber reinforced poly-mer composites can be observed. When mechanically loaded for in-plane rotation and elongation, these fab-rics initially oriented at ±45◦with respect to the loading
Figure 1 Schematic load (P) vs. deformation (ε) curve, or compliance curve, of the media. E1and E2are elastic moduli characteristic of the elastin and collagen phases, respectively. The schematics represent con-ceptually the deformation of both phases under pressure.
direction present a concave shape, or “J” shape, load vs. extension curve [11]. This shape was attributed to a pure intraply shear deformation, or in-plane rotation, of the fabric structure under uniaxial elongation during which the angle between the two roving directions decreases gradually until a critical angle value, or locking angle, is obtained as a result of the complete rotation of the rovings. Once this locking angle is reached, the com-posite rovings progressively undergo extension and the slope of the load vs. extension curve increases rapidly as a result of their high modulus. This behavior has been discussed in the literature [11]. It was shown that the first part of the curve was controlled by the in-plane transverse shear of the fabrics associated to a “trans-verse modulus factor” and the second part of the curve was controlled by the longitudinal deformation of the fabrics associated to a “longitudinal modulus factor” [11]. From this description of the behavior of compos-ite fabrics, it appears that iso-compliant vascular scaf-folds could be obtained from continuous polymer fiber fabrics.
Preliminary results were obtained from thin non-woven mats of polypropylene (PP) fibers using the melt blowing process coupled with an attachment that con-trols the fiber entanglement and fiber distribution in the veil [12]. These mats were then stacked into multi-ply (20) veils and submitted to a heat treatment. Polypropy-lene was selected because of its hemocompatibility and its abundant use for manufacturing commercial suture wires. Polypropylene fibers with 6–14 µm in diameter were produced. The heat treatment, or consolidation step, lasted typically 15 min at a temperature rang-ing between 120 and 150◦C. The nominal thickness of these non-woven fabrics was 0.5 mm. Although PP fibers show an elastic modulus considerably higher than that of the arterial wall, it is possible to control and tai-lor the shape and moduli of the load vs. extension curve of the non-woven fabrics by controlling the inter-fiber bonding and the respective orientation of the individ-ual veils upon stacking. Flat specimens with a nominal 40-mm width were obtained from the non-woven fab-rics and tested in tensile mode at a crosshead speed of
Figure 2 Stress–strain curves of fabric samples prepared (under differ-ent consolidation temperatures) and femoral artery from Ref. 13. Fabric samples with 50% of longitudinal fibers in the stretching direction are represented for two veils (120◦C), four veils (130◦C) and three veils (140◦C and 150◦C). Temperatures refer to consolidation.
12.5 mm/min. These tensile curves are shown in Fig. 2 at different consolidation temperatures. These curves show the characteristic J-shape aspect mentioned, with a weaker first slope (E1) that gradually increases to
a second stiffer slope (E2) at approximately 15% in
deformation. The first slope can be attributed to in-plane rotation of the non-woven fabric structure under uniaxial elongation until the PP fibers are individually stretched, which defines the second slope of the curve at high deformation. These results show that the condi-tions in which the nonwoven fabrics are consolidated, especially the temperature, have an important effect on their shape and first and second slope values. Also in-cluded in Fig. 2 is the curve of in vitro femoral arteries taken from Blondel [13]. Stress-strain curve compari-son between the fabrics and femoral arteries reveal a close similarity in their behavior. They show the same characteristic J-shape aspect and the same moduli and critical deformation. From these similarities in behav-ior, it appears that the first part of the curve in the non-woven fabrics, during which alignment of fabric fibers occurs, corresponds to the elastin fiber re-orientation, while the second part of the curve during which PP fibers are progressively stretched corresponds to colla-gen fiber deformation.
It can be concluded that the behavior of an arte-rial wall can be mimicked using consolidated stacks of non-woven veils of polypropylene. The number of veils, their stacking angle, and their consolidation con-ditions can be used to tailor the mechanical properties of the material. Finally, from the rationalization of this J-shape compliance behavior of textile composites fab-rics into two “modulus factors” [11], it will be possible to adjust the properties of the non-woven PP fabrics to match those of the arterial wall.
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Received 14 July
and accepted 19 October 2004
