Fluorinated Biomaterials for Cardiovascular Surgery

CHAPTER 8

Fluorinated Biomaterials for Cardiovascular Surgery Charles Baquey,1,
Marie-Christine Durrieu,1 and Robert G. Guidoin2 1

INSERM, U577, Bordeaux, F-33076 France; Univ. Victor Se´galen Bordeaux 2, F-33076 France 146, rue Le´o Saignat, 33076 Bordeaux Ce`dex, France 2 Department of Surgery (Biomaterials), Faculte´ de Me´decine, Pavillon Ferdinand-Vandry, Universite´ Laval, Que´bec G1K7P4, Canada Contents 1. Introduction 2. Blood–vessel wall relationships (interactions of flowing blood with the vessel/vascular prosthesis wall) 2.1. Role of the surface free energy or surface tension 2.2. Role of electrical parameters 2.3. Scenario for blood–material interactions 2.4. Role of dynamic factors 2.5. Role of the morphology 3. Requirements for a cardiovascular biomaterial 4. From polytetrafluoroethylene to microporous teflon-based vascular prostheses 4.1. State of the art related to vessel repair or replacement: Evolution and role of PTFE 4.2. How to improve the functional patency of ePTFE-based prostheses? 4.3. Chemical modifications of fluorinated polymers: A way to the improvement of their haemocompatibility 4.3.1. PTFE case 4.3.2. PVDF case 4.3.3. P(VDF-HFP) case 5. Conclusions References Note from the Editors

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Abstract The biomedical area did not miss the tremendous development of macromolecular materials and their applications, which has followed the rapid expansion of petrochemistry during the twentieth century. In fact, these materials are widely used to design a lot of


Corresponding author.; Email: [email protected]

FLUORINE AND HEALTH A. Tressaud & G. Haufe (Editors) DOI: 10.1016/B978-0-444-53086-8.00008-4

# 2008 Elsevier B.V. All rights reserved.

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medical devices that are at physicians’ disposal to care patients. Some of these devices, commonly used in medical care units, such as syringes or tubings, are disposable, whereas others, such as joint prostheses, vascular substitutes, artificial lens, suture thread, are implanted in the body for an unlimited time, hopefully. Considering the cardiovascular area only, the introduction of polymeric materials induced outstanding advances, as they allowed, for instance, the design of arterial substitutes from woven, knitted and non-woven synthetic fabrics, thus opening a new therapeutic option for the replacement of diseased arterial segments beside the grafting of an autologous venous segment, the availability of which is necessarily limited. The polyethylene terephthalate (PET) takes a prominent place among the materials which are used for this application, but it meets the strong competition of the expanded polytetrafluoroethylene (ePTFE). Known for its chemical inertia and great thermal stability, this polymer, which is not easily processed, has many assets (biostability, heat sterilisable) in comparison with its main competitor, the PET, even if like the latter, it does not satisfy all the requirements a material used for the making of arterial substitutes has to comply with. However, different approaches are proposed to improve the situation, such as surface processings specifically designed to increase the polytetrafluoroethylene (PTFE) haemocompatibility on the one hand, or the choice of other fluorinated polymers on the other.

1. INTRODUCTION Because of inherited lesions or of an evolutive disease of their wall, arteries may become unable to ensure an adequate transport of the blood to organs and tissues, which may create a risk of myocardium infarctus or aneurysm rupture followed by a fatal haemorrhage for patients concerned by such (or suffering such) vascular diseases. After Gluck’s pioneering works [1], which paved the way to the modern vascular surgery concepts, the replacement of a carotid segment by an autologous venous graft was no longer a dream, and these patients could be surged to repair or replace their diseased arteries. In this respect, vascular surgeons have at their disposal, various devices which do not satisfy all the requirements which are listed below, but which may compensate some of the functional deficiencies their patients encounter. Following Gluck’s works, various materials including metals and glass have been used for the design and making of vascular substitutes, but the first clinical studies related to the use of flexible prostheses made from a non-biological material were reported only after 1952 by Woorhees et al. [2]. Taking into account what is known about the flowing blood–artery wall relationship, different criteria which have to be satisfied by a blood-contacting material, and more specifically by an arterial substitute, can be identified. This rationally designed strategy, mainly documented by studies carried out during the seventies and later on, was not given the priority as surgeons chose 20 years earlier, pragmatic solutions based on the use of polymeric materials among which polyethylene terephthalate (PET) and polytetrafluoroethylene (PTFE) a little later took leading positions.

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2. BLOOD–VESSEL WALL RELATIONSHIPS (INTERACTIONS OF FLOWING BLOOD WITH THE VESSEL/VASCULAR PROSTHESIS WALL) Interactions of blood, or more precisely of plasma macromolecules and blood cells with the vessel wall, should the latter be natural or artificial, depend upon several parameters: the mechanical properties of the vascular conduit, on the one hand, the morphology and the physical and chemical characteristics of the blood-contacting surface on the other.

2.1. Role of the surface free energy or surface tension Surface tension, or surface free energy, is a parameter which corresponds to the residual binding capacity of a material surface, that is, the binding capacity of atoms or groups of atoms which constitute the border surface of the material of interest. This binding capacity may be uniform, as is the case for non-oxidised metals, but most of the time it is the resultant of several components that are respectively related to the various types of atoms or atom groups present on this surface including ionic sites, hydrophobic sites, polar sites or hydrogen atom donors or acceptors. Accordingly, most of the material surfaces appear as mosaic-like structures. The nature of the potential sites of interaction and the microtopography determine the interaction phenomena of materials with biological media and especially with blood.

2.2. Role of electrical parameters Blood cells and vessels walls are negatively charged, the corresponding isoelectric point lying at a pH between 4.8 and 5. Sawyer and Srinivasan [3] measured a voltage difference between endothelial layer and blood. Negatively charged surfaces give rise in the presence of an electrolyte solution to a double electric layer responsible for the recorded potential value C0 (Fig. 1a). As the distance from the wall increases, C0 decreases linearly down to a value equal to C0/2.3 according to the Stern theory. The corresponding distance (l) to the wall is known as the Debye length. At distances greater than l, the potential decreases exponentially to zero, the decaying potential being characterised by the so-called zeta (z) potential (Fig. 1b). The distance from the wall, corresponding to z, defines the shear plane and depends upon the flow conditions. For a given electrolyte solution (i.e. the nature and concentrations of ions in blood are known), z increases linearly with the flow rate [4]. Since the vessel wall is negatively charged, it repels the negatively charged platelets, and it helps in preventing thrombogenic phenomena. Experimentally, Sawyer et al. [5] demonstrated that on positively charged surfaces [>200 mV/normal

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+ + + + + + + + + + +

− − − − −

− − −

− −

− − −

−

Potential Y

(b) Y0

Y0 /2.3 z 0

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Fig. 1. Variation of the electric potential near a surface in the presence of an electrolyte solution. (a) Electrical double layer at the surface of a solid positively charged, in contact with an electrolyte solution. (b) The variation of the electrical potential when the measurement is made at an increasing distance from the surface, and when the liquid phase is mobile at a given flow rate. The zeta potential (z) can be calculated from the streaming potential, which can be measured according to the method described by Thubikar et al. [4].

hydrogen electrode (NHE)] thrombosis occurred systematically, while thrombosis never occurred on surfaces with a negative potential (<0 mV/NHE). However, surface potential cannot be the unique criterion of non-thrombogenicity. The superficial distribution of the charged site plays an important role as far as plasma protein adsorption is concerned and, upon their adsorption, these biomolecules may trigger the coagulation cascade in spite of exposing a net negative charge to the bloodstream.

2.3. Scenario for blood–material interactions According to the above developed considerations, surfaces can be described as a mosaic-like structure of which individual elements contribute locally and specifically to their residual binding capacity and their electrical potential. The individual characteristics of these elements and their microtopography determine, via a first series of interactions with water molecules and small solutes, which plasma

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proteins adsorb and how they do it, when blood comes into contact which these surfaces, according to the following scenario. At first, and according to their surface tension, materials show an affinity for water which corresponds to their wettability. Polar materials and materials bearing ionic sites, whether these properties are intrinsic or brought about by the contacting medium through adsorption phenomena of ions, are obviously the most wettable. Hydrophobic materials do not interact with water, but instead promote its intermolecular organisation, as in the case of oil or parafins. The contribution of water to the entropy of the system is thus decreased. This first step may occur simultaneously with (or may be immediately followed by) the interaction of free ions in solution on the one hand, which interact with ionic sites present on the material surface for previously mentioned reasons, and the interaction of small solutes in solution on the other. Obviously, ions and small solutes contribute to a local alteration of the water organisation in the interfacial region. The second step concerns proteins and other biological macromolecules, which will adsorb at the blood–material interface. These adsorption phenomena are controlled by the nature and distribution of available binding sites as determined by the type of material and the state of its surface after the previous step. Biological macromolecules are themselves characterised by a superficial distribution of binding sites when they are free in solution. If there is a good fit between this distribution and the one exposed by the material surface, macromolecules may adsorb without any conformational change; among available macromolecules, those for which this fit is the best have the greatest probability to adsorb. However, adsorption may occur together with a conformational change if such behaviour is thermodynamically favoured [6]. Furthermore, these adsorption phenomena may give macromolecules an opportunity to be activated, that is, zymogens will give active enzymes. In this way, the so-called coagulation contact phase starts. The third and final step involves blood cells, and more particularly platelets and polymorphonuclear neutrophil leukocyte also known as granulocytes. According to the material surface state after the various events which have occurred previously, cells may or may not adhere to the material; for platelets, adhesion may be the initiating step for their activation and aggregation, since activation of platelets involves the release by their cytoplasmic granules of active promoters for thrombogenesis and platelet aggregation. Cell adhesion is not only due to physical and chemical phenomena as for protein adsorption, but also implies biochemical mechanisms involving cell membrane glycoproteins. C5 to C9 complement{ components may bind to platelets and amplify platelet release and aggregation already induced by thrombin. It may be noticed that platelet activation may also result from contact with air, which cannot occur under

{ The so-called complement system consists in a set of 21 proteins aimed at contributing to the defence of the organism against foreign bodies or microorganisms or even abnormal cells.

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implantation conditions, but which is possible during a session of extracorporeal circulation. Similarly, neutrophil adhesion arises from an increase of their adhesiveness due to the effect of the anaphylatoxin C5a on the expression of adhesines at the membrane of the neutrophils; C5a is generated by the complement activation processes, which may occur during the plasma protein–material interaction step.

2.4. Role of dynamic factors Surface physics and chemistry play a major role as far as blood–material interactions are concerned. However, the observed processes depend as much on the local flow patterns of the blood near the material surface as on these surface characteristics. It has been shown that plasma protein adsorption, which occurs as soon as blood comes into contact with a given surface, is clearly influenced by the local shear stress [7]. As discussed by Leonard [8], the concentrations of activated coagulation proteins in the interfacial layer (i.e. the layer sitting between the main stream in a vessel and the wall) are strongly influenced by flow factors, and the net concentrations are determined both by biochemical reactions producing activated coagulation factors and by convective–diffusive factors that serve to moderate the concentrations of inactivated and activated coagulation proteins. Furthermore, the presence of high concentrations of activated coagulation proteins may favour the activation of platelets and leukocytes. These two blood cell types are able to undergo activation via collisions, which are favoured by turbulent flow or particular conditions such as those created by vortices. Platelet and fibrin deposition are strongly influenced by plasma and cellular factors, by properties of the vessel wall and by flow, as shown experimentally with the model of the everted deendothelialised rabbit aorta, used by Baumgartner [9]. Platelet attachment to subendothelium{ is determined predominantly by physical factors controlling the rate of platelet transport to the subendothelium at low shear rates (800 s1). In addition, platelet deposition was found to be highly dependent on the concentration of red cells, an effect attributed in part to the fact that red cells, by increasing the radial movement of platelets, enhance their diffusivity by several orders of magnitude compared to that theoretically predicted and experimentally measured in platelet-rich plasma. Weiss et al. [10] have shown that platelet deposition is at a minimum at a shear rate of 50 s1, while fibrin deposition on subendothelium from non-anticoagulated blood is at a maximum at the same shear rate. At low shear rates (250 s1), fibrin { The inner surface of vessel is composed by a monolayer of adjacent endothelial cells. This layer is called endothelium and the underlying tissue is called subendothelium.

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deposition is independent of platelet density and integrity. At a higher shear rate (650 s1), platelets must possess all their physiological properties to promote fibrin deposition. At much higher shear rates (2600 s1), fibrin deposition decreases despite increasing platelet deposition. It will be later proposed that an ideal blood-contacting surface would be one on which endothelial cells could adhere and grow normally to create a complete endothelium endowed with the physiological functions of such tissue. But the correct expression of such physiological functions requires normal blood flow conditions. And the correlation of thrombosis with regions of disturbed flow suggests that shear stress may alter the production of endothelial cell-derived products and directly affect endothelial cell function. Secretion of tissue plasminogen activator by cultured endothelial cells increases within an hour after exposure to arterial levels of shear stress [11], while secretion of the related inhibitor (plasminogen inhibitor-1) by the same cells is unaffected by shear forces over the physiological range. Endothelial cells produce vasoactive substances, which modulate the permeability of vessel walls; accordingly, shear stresses may indirectly affect this wall parameter by influencing the secretory activity of endothelial cells. Prostacyclin, a potent inhibitor of platelet aggregation, is derived from metabolisation of arachidonic acid by endothelial cells, and shear stress increases its production rate [12]. It is postulated that this effect is due to perturbations of the permeability of the plasma membrane changing the cytosolic Ca2þ content and leading to an increase in phospholipase C activity (through the by-passing of the receptor requirement), which contributes to a higher production of arachidonic metabolites. Shear stresses may also affect the construction and secretion by human endothelial cells of von Willebrand Factor polymeric forms that are involved together with fibronectin and fibrinogen during shear-induced platelet aggregation. In addition, shear stresses may play a role in the mechanism of platelet aggregation, probably through an increase in the readily available amounts of ADP, should the latter be due to an increased lysis of platelets, or to a greater release of their granules content.

2.5. Role of the morphology Materials may be compact or porous. In the case of compact materials, smooth surfaces are preferred, as blood cells will encounter fewer opportunities to be injured by asperities or morphological singularities when blood flows in contact with such surfaces. The chance of cell trauma of mechanical origin depends obviously upon the size of these morphological irregularities and upon local flow conditions (shear stress, tubulences, vortices, etc.). Wurzinger and SchmidSchonbein [13] have shown that the number of platelets adhering to PVC

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surfaces could be multiplied by a factor of 3 with surfaces featuring peaks or valleys with averaged heights or depths around 9 mm compared to polished surfaces, knowing that the diameter of platelets sits around 1.5 mm. Porous materials may have a closed porosity, which lowers their density and can be varied to optimise their mechanical properties, as this is carried out for some kinds of vascular grafts. For such materials, surface morphology must fulfil same requirements as those fulfilled by compact materials. Materials with an open porosity are not generally directly exposed to flowing blood for long. A widespread example of such materials is found with knitted or woven PET-based vascular grafts. These grafts may be immersed into blood to allow the latter to invade the fabric mesh where it coagulates, the resulting matrix acting as a new lining offered to the flowing blood.

3. REQUIREMENTS FOR A CARDIOVASCULAR BIOMATERIAL The above considerations show how it is important for the design of a vascular substitute to pay attention to its mechanical properties, to get similarity of dynamic behaviour between this substitute and the vessel it is to partly replace. The bloodstream that will flow through this artificial conduit must not suffer any alteration able to induce any local turbulences or high wall shear rate. This requirement may be satisfied insofar as the mechanical compliancy of the prosthesis (see Table 1 for definition) is equivalent to that of the natural artery it will replace. In fact, the arterial blood flow is pulsatile and characterised by a pressure wave which should propagate itself across any section of the arterial tree without encountering any sudden change of mechanical impedancy (potentially responsible for energy consuming reflected waves). The viscoelastic properties of their wall, featuring a composite structure, allow natural arteries to increase

Table 1. Comparison of the respective evolution of the compliancies of a natural artery on the one hand and of an expanded polytetrafluoroethylene prosthesis on the other hand when the blood pressure increases Blood pressure (mmHg)

Femoral artery Expanded polytetrafluoroethylene

60

80

100

120

6.5 0.9

4.7 0.85

4.1 0.78

3.8 0.8

Compliancy: Measures the ability of an artery to increase its section (and obviously the volume of blood, which can be contained in a given segment of this artery) upon an increase of TransMural Pressure (PTM) compliancy is expressed as ((DV/V)  100)/DP as % per mmg Hg.

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their diameter according to a non-linear pattern, as the blood pressure increases; the greater this pressure, the smaller the relative increase of the diameter (see Table 1). It is expected that the laminar blood flow regimen will not turn to a turbulent one, when it crosses over an anastomosis between a natural artery and an arterial prosthesis (or between the latter and again the natural artery), if the prosthesis has a compliancy similar to that of the natural artery. The turbulent regimen is undesirable as it increases the probability of collision between blood cells and causes great variations of the wall shear rate, the consequences of which are detrimental in several respects: intimal lesions, platelet activation, peri-anastomotic hyperplasia and false aneurysms formation. Whichever is the targeted application several fundamental requirements must be satisfied. Devices designed for the replacement of arteries, and then the materials used for the making of these devices, must be mechanically adapted to suffer for many years stresses resulting from the pulsatile blood pressure. They must feature as well, an adequate level of haemocompatibility, this criterion being easier to satisfy as vessels to be replaced have a larger diameter. Lastly, they must be biocompatible, which means that they are able to promote adhesion and healing of surrounding tissues without inducing any heavy inflammatory process or fibrosis or any other undesirable process. As a matter of fact, chemically inert materials, such as polyesters like PET on the one hand, or PTFE on the other, drew the interest of specialists, and were used as woven or knitted fabrics (for PET and PTFE initially), or as non-woven fabrics for PTFE (later on) to make tubular conduits; the porous wall of the latter was supposed to favour the above mentioned healing process. In fact, the wall porosity allows a transmural ingrowth of blood capillaries, the latter providing a potential source of cells for the endoluminal endothelialisation of the artificial substitute on the one hand, and oxygen and nutriments to the neointima on the other as does the vaso vasorum capillary network through the wall of native vessels. In addition, this parietal microvasculature helps to strengthen the host’s defences against infectious processes, of which the risk is objectively increased by the implantation of a prosthesis, as bacteria easily adhere onto the surface of the latter. However, a compromise must be established between the above mentioned advantages brought by vessel substitutes with a porous wall and the blood loss suffered by the patient as soon as his blood is allowed to flow through such substitutes. In fact, the blood stops leaking rather rapidly as it clots when it comes into contact with the artificial surface of the vascular prosthesis. Nevertheless, and as a consequence of this blood loss, patients may need a blood transfusion, which may expose them to specific risks. That is the reason why vascular prostheses makers proposed in the early eighties, devices with an impervious wall. Such devices were prepared by soaking classical woven or knitted polyester prostheses into an albumin or collagen solution, and by a further cross-linking of the absorbed protein [14]. In vivo, the resulting proteinaceous reticulum is

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Fig. 2.

A type of surgical vascular suture.

progressively resorbed, and a few days after implantation, the primary porosity of the prostheses is again available for an eventual ingrowth of blood capillaries. Last but not least requirement, artificial arterial substitutes must be able to be connected to the host’s vessels using sutures, the only reliable mean surgeons trust (Fig. 2). The connections performed must be stable and blood tight for as long as the prosthesis will remain patent. This last condition has supported the interest of surgeons for vascular prostheses made of woven or non-woven and knitted synthetic fabrics. The following sections give an outline of the different steps that line the evolution of materials for the cardiovascular system, and present some prospective solutions that have been proposed and supposed to improve the performances of these materials.

4. FROM POLYTETRAFLUOROETHYLENE TO MICROPOROUS TEFLON-BASED VASCULAR PROSTHESES Among these materials, two polymers the PET and the PTFE, which was discovered in 1938 by Roy J. Plunkett (after Peska et al. [15]), hold a predominant place, and obviously our comments will mainly concern the second one. The perfluorination of ethylene leads to the starting vinyl monomer for the synthesis of the polyethylene perfluorinated homologue; according to the propylene content of the starting hydrocarbon, the resulting fluorinated polymer may include a more or less great proportion of perfluorinated statistic copolymers of ethylene and propylene; the so-called Teflon FEP corresponds to one of such blends. Because of the high energy of the C–F bond, PTFE is extremely stable; its high molecular weight (>106 g/mole) renders it insoluble in all usual organic solvents, and resistant to relatively high temperatures, as it can suffer permanently as high temperatures as 260 C. Because of the strong electronegativity of fluorine, PTFE is not

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flammable, and demonstrates a high chemical inertia. Moreover, its surface free energy (18 ergs/cm2) is the lowest in comparison to that of usual polymeric materials; in fact, PTFE is the hydrophobic reference material, its hydrophobicity being likely responsible for conformational changes (with already mentioned biological consequences) proteins suffer when they adsorb onto its surface. Its high melting temperature (>320 C) and the high viscosity of the melted polymer prevent its transformation according to procedures used with classical thermoplastics. Accordingly, a specific transformation protocol, derived from the powder metallurgy technology, has been developed to produce flexible tubings fitting with surgeons’ wishes. According to this protocol, PTFE pellets are mixed with naphthalene, which plays a role of lubricant during the high temperature extrusion of this blend into tubes. The latter are modified by a process including several cycles of heating and mechanical stretching. This process, which is carried out at a temperature close to the PTFE melting point and favours the coalescence of PTFE crystallites, as well as induces the elimination of residual naphthalene, strongly alters the structure of the wall and create a macroporous, non-woven fabric with a high percentage of its volume composed of transmural void spaces. The structure that is formed is called expanded polytetrafluoroethylene (ePTFE), consists of layers of short, interlocking, circumferentially oriented, solid bands of the polymers (nodes), which are linked together by numerous, fine, filamentous bands (fibrils) that are longitudinally oriented, respect to the graft, and perpendicular to the nodes (Fig. 3A,B). Grafts have been investigated with internodal distances ranging from 5 to 90 mm in an effort to optimise haemocompatibility and endothelial cell attachment; however, the grafts in current clinical use, all have average internodal spacing of about 30 mm. These grafts are easily handled by surgeons and can be sutured securely; however, they have two drawbacks: they are prone to kink under severe bending and to collapse under the compressive action of surrounding tissues. To prevent these bad phenomena, companies have proposed to stake the graft wall with a propylene rib, but the abrasive effect of the latter on the material wall during the prostheses use induced potentially more severe risks for the patients.

4.1. State of the art related to vessel repair or replacement: Evolution and role of PTFE Like PET, which further became more and more popular under various trade names (Dacron in USA, Rhodergon in France, Dallon in Russia and European Eastern countries), PTFE was initially used as thread, which could be woven or knitted to obtain vascular prostheses. These resulting woven devices, in which two sets of threads, respectively called weft and warp, cross each other perpendicularly, are known to easily fray near anastomoses. For knitted devices, in

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Fig. 3. Microstructure of expanded polytetrafluoroethylene as shown by scanning electron microscopy with different internodular species (a) 30 mm; (b) 90 mm; (c) 12 mm. (N: Nodules, F: Fibrilles)

Fig. 4. Vascular prosthesis made of woven polytetrafluoroethylene. The photograph shows the great ability of the fabric to fray.

which these two sets of threads may not only interlace but also lock up with each other, fraying in less important, but still occurs due to the PTFE fibre characteristics. The ends of the primary devices as proposed by Edwards [16] (after Harisson experiments [17]) frayed rather easily (Fig. 4), especially when they were beveledged. However, these devices showed rather good clinical performances in spite of an uncertain healing. Ruptures might occur near the anastomoses when

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silk was used to realise the sutures, as silk is partly biodegradable, while Teflon fibres kept unaltered for several decades after implantation [18]. Nevertheless, woven PTFE arterial prostheses were no longer used, as surgeons preferred to use prostheses made from woven or knitted PET yarns. Later on, these polyester-based prostheses met the competition with prostheses made from microporous Teflon (Fig. 5) or ePTFE, better known under various trade names such as Gore-Tex, IMPRA, Vitaflon, etc. Microporous Teflon was first used by Ben Eiseman in 1972 [19] for the making of membranes fitting oxygenators used for extracorporeal circulatory assistance. As several benefits seemed to be related to the use of this material (less damaged blood cells, no or little protein depletion, no or few clot deposition, neither embol release), he tried with his team to extend its applications. They carried out a series of experiments with piglets, such as the implantation of microporous Teflon conduits as substitutes for the portal vein, the inferior vena cava, the external iliac vein, and obtained a permeability rate equal to 80%, 2 months after implantation, taking into account the data from 27 experiments. In a pancreatectomised patient (pancreatectomy being a consequence of a cancer), 32 months after the implantation of such an artificial graft as a substitute for the portal vein, these graft was still permeable. Such success encouraged major companies, such as W.L. Gore and IMPRA, and vascular surgeons to develop the use of such grafts as arterial substitutes, and several teams carried out experimental studies which show them performant for various implantation sites. The first series of femoro-popliteal shunts, including 15 patients suffering a severe lower limb arterite (level III, even IV) and for whom the saphenous vein could not be used, was reported by Campbell in 1976 [20]; the patency rate at 8 months was 87%. These ePTFE prostheses made surgeons very enthusiastic and at least one of them, Veith [21], argued in their favour through an impressive

Fig. 5. Vascular prosthesis made of micoporous polytetrafluoroethylene, also known as expanded polytetrafluoroethylene.

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Fig. 6. Tubular conduit made of expanded polytetrafluoroethylene wrapped into a thin sheath, which is also made of polytetrafluoroethylene; the photograph shows the ability of the sheath to detach from the main wall surface.

clinical experience. Few cardiac surgeons attempted to use these prostheses for aortocoronary bypass, but they gave up very soon with this practice. The analysis of microporous Teflon-based arterial prostheses, which were retrieved along autopsies or along revision surgeries, confirmed the excellent biostability of these devices insofar as they had been carefully implanted. Few ruptures or tears were observed but had been likely caused either by a bad handling or by fibrosis expansion, the latter being more frequent for nonsheathed devices (Formichi et al. [22,23]). Taking into account the lack of stability of the prosthesis wall, the two major makers made tremendous efforts to improve this situation. W.L. Gore chose to fit their prostheses with a very thin external sheath, made also of microporous PTFE but almost impervious (Fig. 6).

4.2. How to improve the functional patency of ePTFE-based prostheses? These prostheses need to be improved in at least two respects: their haemocompatibility that must be increased if they are to be used for the replacement of small diameter artery (f<4.5 mm) on the one hand, and their dynamic mechanical properties on the other. To resolve the first issue, two strategies that in fact do not apply specifically to ePTFE-based prostheses have been proposed. According to a first strategy, the triggering of the blood coagulation cascade, which generates thrombin as soon as blood comes into contact with the artificial material, cannot be avoided, and the latter must be endowed along appropriate chemical modifications, with the ability to catalyse the inhibition of thrombin by its naturally circulating inhibitor, that is, the so-called antithrombin. Among the proposed modifications, one can distinguish the covalent binding of either heparin [24],

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a specific glycosaminoglycan of which the catalytic effect on the thrombin– antithrombin reaction is well known, or synthetic polymers that are functionalised along a further stage to endow them with heparin-like properties [25]. A second strategy aims at avoiding the triggering of the blood coagulation cascade by preventing in a non-specific way, the adsorption of the so-called contact phase proteins and their conformational changes upon adsorption on a surface as hydrophobic as PTFEs, these conformational changes being responsible for their activation. Pre-adsorption of human serum albumin has been shown empirically to be somewhat efficient in this respect. According to a more rational approach, irreversible adsorption of proteins in general, and more particularly of proteins belonging to the coagulation system, can be prevented by the presence at the surface of the PTFE of a hydrophilic layer of which the main constituent may be polyethyleneglycol molecules. The latter can be bound to the PTFE insofar as it has been previously adequately functionalised; a plasma treatment, as described in a further section, allows such a functionalisation [26]. However, the expected best way to prevent blood to clot upon its contact with prostheses wall is to coat the latter with endothelial cells, which are used to regulate the blood–vessel wall relationship. To carry out this coating, the material surface has to be endowed with pro-adhesive properties towards endothelial cells; such properties may be brought by peptidic ligands able to be specifically recognised by specialised membrane receptors (integrins) sitting at the surface of the cells. These ligands are characterised by the presence in their sequence of the RGD (arginin-glycin-aspartic acid) triade [27], the latter being also included in the sequence of extracellular matrix proteins (collagen, fibronectin, vibronectin, etc.). Thus, such tripeptides may be also brought onto a prosthetic surface by coating the latter with an aqueous solution of collagen or of any other extracellular matrix protein. To control their availability and their superficial distribution density, proadhesive synthetic peptides may be used, as reported in a subsequent section. Endothelial cells anticoagulant properties are partly due to their ability to express at their membrane, thrombomodulin (TM) a glycoprotein, which tightly binds thrombin. So TM works as a direct anticoagulant by inhibiting the thrombin cleavage of fibrinogen, the first step of the fibrin clot formation; more importantly, TM is the essential cofactor that promotes thrombin cleavage of protein C; the resultant activated protein C inactivates factors Va and VIIIa and thus shuts down coagulation and haemostosis. This explains the choice made by Vasilets et al. [28] who proposed to immobilise TM onto PTFE previously grafted with acrylic acid (AA) through a microwave CO2 plasma initiated process. Considering the second requirement, that is, good dynamic mechanical properties ePTFE-based arterial prostheses must satisfy to improve their functional patency, there are not so many ways to reach this goal. Unfortunately, and due to the intrinsic mechanical characteristics of PTFE, the mechanical behaviour of these prostheses cannot be easily modified. However, different ways have been explored; on the one hand, an optimisation of the design of the prostheses

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has been attempted; on the other, the replacement of PTFE by other fluorinated polymers with a lower Young’s modulus has been envisaged. As far as the design of the prostheses is concerned, the replacement of strictly cylindrical conduits, by tapered ones that better conform to the anatomic realty, has been proposed and deeply documented by theoretical and experimental studies [29]. ePTFE-based tapered arterial prostheses have been commercialised, but did not appear to bring any clinical benefit. The design of prostheses with an oval section, instead of a circular one, offers a way to get a gain in compliancy as the section area of such prostheses increases via a simple change of its shape (oval ! circular) as blood pressure increases, without any change of its circumference. However, to our knowledge, no arterial prosthesis designed according to this concept has ever been commercialised. The poor ability of ePTFE to reversibly stretch is not compatible with the making of arterial substitutes with a good compliancy; thus, specialists considered other polymeric materials combining the chemical inertia and the biostability of PTFE with a mechanical behaviour similar to that of elastomers, and their interest moved towards polyvinylidene difluoride (PVDF) and even more towards copolymers of vinylidenedifluoride and hexafluropropylene (HFP), which may have pseudo-elastomeric properties. Until now, only PVDF has been processed into yarns [while yarns of P(VDF-HFP) copolymers are not commercially available] from which tubular conduits, prefigurating vascular prostheses, have been woven (see Section 4.3). Even if the mechanical behaviour of these woven PVDF-based conduits, when they are experimentally exposed to a pulsatile flow with characteristics analogous to that of blood flow, is better than that of ePTFEbased prostheses or PET-based ones, it is far from conforming specialists’ expectations. Nevertheless, the faisability of the chemical modification of PVDF to improve its haemocompatibility, according to the previously described principles, has been investigated. A further step will concern the modification of P(VDF-HFP) copolymers (hopefully, the same procedure as the one used with PVDF will be applicable), as soon as corresponding yarns from which tubular conduits demonstrating the expected mechanical behaviour could be woven, will be available.

4.3. Chemical modifications of fluorinated polymers: A way to the improvement of their haemocompatibility This section will successively deal with the treatment of PTFE, the treatment of PVDF and the treatment of P(VDF-HFP) copolymers.

4.3.1. PTFE case Because of its high chemical inertia, the chemical modification of PTFE may appear as a challenge. However, although this polymer resists to acidic or alkaline attacks,

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it is vulnerable to the action of ionising radiations and to gaseous plasmas. Unfortunately, when PTFE is exposed to g-rays or to accelerated electrons beams, it degrades even if the adsorbed doses (a few kilograys) are relatively low; its molecular weight and its tensile strength decrease dramatically, as a lot of structural bonds are broken in the whole bulk of the material. Plasma treatment offers the opportunity to break bonds that belong only to the very first layers of macromolecules sitting at and under the surface of the material, without any noticeable bulky effect. Among others authors, Vasilets et al. [28] already quoted and Baquey et al. [26] demonstrated the interest of such an approach to introduce some functionality at the surface of ePTFE. After their paper, Baquey et al. exposed ePTFE samples to a Radio Frequency Glow Discharge fed with an argon/oxygen mixture (flow rate ratio 1:1) in a classical reactor. The latter was fitted with two internal aluminium electrodes, one being grounded and the other powered by a 13.56 MHz generator and acting as a sample holder. The pressure was set at 100 mTorr, and the treatments were driven at 100 W for 1 min. Then the samples were exposed to air atmosphere for 1 h before further grafting by AA via their immersion in an aqueous Acrylic Acid (AA) solution (25%, v/v) contained in glass flasks; following vigorous degassing, the flasks were sealed and kept for 5 h in a water bath at 65 C. The AAgrafted samples were taken out of the flasks, and appropriately washed to remove any adsorbed homopolymers. Along a further step, the carboxylic groups available at the surface of the AA-grafted samples were exploited for the coupling of bNH2PEG, which is intended to be used as an anchor arm between the surface and peptidic molecules of interest via amide bonds. In this respect, Baquey et al. [26] chose a tetrapeptide, containing the RGD sequence and being, for that reason, a potential ligand for integrins (adhesion molecules), which sits at the membrane of cells, and particularly, of endothelial cells. The different steps of the procedure were monitored by X-ray photoelectron spectroscopy (XPS) quantitative analysis. It appeared that after plasma treatment (according to the above described protocol here) the oxygen content of the material surface was significantly increased, this being possibly attributed to the formation of peroxides. The latter were supposed to give rise upon heating, to peroxyl radicals able to induce the grafting of AA. In fact, the oxygen content of the material surface further increased after AA-grafting, and the presence of carboxylic groups was confirmed by a colorimetric assay based on the reactivity of these groups towards toluidine blue. So was confirmed the coupling of the AA-grafted surface with bNH2PEG, as the presence of nitrogen could be detected as significant by XPS. The final coupling of the RGDC tetrapeptide (Arg-Gly-Asp-Cys) was confirmed, as well as the presence of sulfur due to its cysteine amino acid component, and could be detected by XPS also. As far as their biological properties are concerned, the treated ePTFE samples showed their ability to better induce the adhesion of endothelial cells than does the pristine material. The number of cells adhered to the material

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after 2 h of culture in the absence of serum were four times larger for the treated ePTFE than for the pristine material; the mechanism of this adhesion was specific as the phenomena could be inhibited by the introduction of soluble RGDC peptides in the immersion medium of the samples containing the cells.

4.3.2. PVDF case PVDF is mainly obtained by radical polymerisation of 1,1-difluoroethylene; ‘head to tail’ is the preferred mode of linking between the monomer units, but according to the polymerisation conditions, ‘head to head’ or ‘tail to tail’ links may appear. The inversion percentage, which depends upon the polymerisation temperature (3.5% at 20 C, around 6% at 140 C), can be quantified by 19F or 13C NMR spectroscopy [30] or FTIR spectroscopy [31], and affects the crystallinity of the polymer and its physical properties. The latter have been extensively summarised by Lovinger [30]. Upon recrystallisation from the melted state, PVDF features a spherulitic structure with a crystalline phase representing 50% of the whole material [32]. Four different crystalline phases (a, b, g, d) may be identified, but the a phase is the most common as it is the most stable from a thermodynamic point of view. Its helical structure is composed of two antiparallel chains. The other phases may be obtained, as shown by the conversion diagram (Fig. 7), by applying a mechanical or thermal stress or an electrical polarisation. The b phase owns ferroelectric, piezoelectric and pyroelectric properties. The interest for this polymer came as already said, from its relatively high chemical inertia, although lower than that of PTFE, and its lower Young’s modulus (1.45 GPa) in comparison to that of PET (10.56 GPa) and PTFE. In addition, the availability of multifilament yarns obtained by extrusion from bulky PVDF offers the possibility to weave or to knit, as it is the case for PET yarns, tubular conduits usable as vascular prostheses (Fig. 8). Moreover, threads made of PVDF have long been actually used in surgery for suturing purposes [33] and have demonstrated a clinical biocompatibility [34] equivalent to that of other non-biodegradable polymeric materials used for this application. In parallel, the possibility to endow this polymer with ‘heparin-like’ properties was demonstrated [35–37]. Briefly, the polymer available as 25-mm-thick films was grafted with polystyrene along a first step including its irradiation either with g-rays delivered by a 137Cesium source, or with swift heavy ions (SHI; mainly Argon) available from the GANIL (Grand Acce´le´rateur National d’Ions Lourds) facility in Caen, France. Different grafting yields (Y %) were obtained according to the type of radiation that was used, the adsorbed dose and the time of exposure to the vinylic monomer, that is, styrene; then the phenyl rings of the grafted polystyrene were chlorosulfonated and the chlorosulfonate groups reacted with the aspartic acid dimethylester; lastly, an alkaline treatment regenerated the carboxylic groups of the aspartic acid residues and generated sulfonate functions from the unreacted chlorosulfonate.

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High T High P

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DME, DMA or DMSO

Solution Cyclohexanone

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Fig. 7. Polyvinylidene difluoride conversion diagram.

Knowing the deleterious effects of ionising radiations on PTFE, much attention was paid on their effects on PVDF although the latter had been told to behave well under irradiation. As an example, PVDF multifilament yarns can be g irradiated up to the absorption of 80 kGy without any effect on their Young’s modulus [38]. However, the structure of the polymer was somewhat modified as its energy to break kept constant and even increased by about 50% while the adsorbed dose was less than 8 kGy but decreased for higher adsorbed doses, down to 1/3 of its maximum value (1/2 of its initial value) when the adsorbed dose reached 81 kGy. To avoid these bulky effects, the interest of

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Fig. 8. Tubular conduit made of woven polyvinylidene difluoride.

Fig. 9.

Scheme of a swift heavy ion latent track.

the irradiation by SHI has been investigated. Instead of being homogeneously distributed in the whole bulk of the irradiated material when g-rays are used, the energy (or the absorbed dose) brought by heavy ions is deposited very locally along so-called latent tracks (Fig. 9) related to their very short range inside the

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material. According to the fluence (i.e. the average number of ions crossing a 1 cm2 section) of the heavy ions beam that is operated, these latent tracks may be either completely separated or overlap with one another. Moreover, due to their high Lineic Energy Transfer (LET), SHI are rapidly stopped; thus, their interactions with the iradiated polymer only concern the very superficial layers of the latter. As a matter of fact, the subsequent grafting of polystyrene can only occur onto the related surface microdomains, where the density of precursors of polymerisation initiators is relatively high. This appears clearly on microphotographs obtained by Scanning Electron Microscopy (SEM) of treated films which show the presence of circular islands (Fig. 10), the diameter of which depends upon the adsorbed dose. In addition, the elemental analysis of these films by Energy-Dispersive X-ray Analysis (EDXA) evidenced a decrease of the relative concentration of fluorine related to an increase of the grafting yield, and this decrease being maximum when the microprobe targets the centre of the above mentioned islands. Fourier Transform IR, transmission and internal reflection spectroscopies have been used to calculate the PS grafting yields at different depths; the observed gradient is higher in the case of g-ray-induced grafting than in that of a SHI one. This might be explained by the fact that the PS growing chains diffusion is accelerated in the latent tracks formed after irradiation with SHI. The existence of a grafting can be linked to the monomer ability to diffuse towards the grafting centres. The subsequent steps of the treatment aimed at endowing the PVDF with heparin-like properties lead as well, to a non-homogenous distribution of the related functional groups; as a matter of fact the hydrophobicity of the material surface depends upon the coordinates of the point around which the wettability is measured. In fact, the surface has acquired a mosaic-like structure; in other words, it consists of two classes of microdomains more or less delimited, respectively characterised by the original hydrophobicity of PVDF for some of them, and by a relative hydrophilicity for the others. According to the concept developed by Okano et al. [39], such surfaces, insofar as they are properly designed in terms of size and specific characteristics of their constitutive microdomains, could express a low thrombogenicity as they could suppress platelet adhesion, and not induce any activation of the coagulation cascade. In spite of these very promising perspectives, there was a lot of work to carry out to apply this concept to the processing of PVDF multifilament yarns. So and considering practical reasons (i.e. the necessity to rapidly obtain tubular conduits which could be experimentally implanted in vivo as arterial substitutes), Marmey et al. [37] decided to initiate the chemical functionalisation of this material through its irradiation with g-rays; even under such conditions, the resulting functions are homogeneously distributed at the surface of the material. To evaluate the effect of the treatment on the haemocompatibility of the PVDF yarns, the so-called thrombin time (TT) of human plasma was measured, whether a given amount of PVDF yarns, either treated or not was, immersed into a plasma aliquot or not. The immersion of pristine PVDF did not change the value

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Fig. 10. Scanning electron microscopy microphotographies of the surface of different polystyrene-grafted polyvinylidene difluoride (PVDF) surfaces; the grafting was performed upon irradiation with Argon ions with various fluences or absorbed doses. (a) PVDF-gAr-PS (Y ¼ 2.5%), D ¼ 3.72 kGy, Ft ¼ 1.1109 ions/cm2; (b) PVDF-gAr-PS (Y ¼ 19%), D ¼ 3.72 kGy, Ft ¼ 1.1  109 ions/cm2 and (c) PVDF-gAr-PS (Y ¼ 12%), D ¼ 37.2 kGy, Ft ¼ 1.1  1010 ions/cm2.

of TT in comparison with the value (18 s) measured for plasma alone, while the immersion of treated PVDF dramatically increased the TT up to values greater than 120 s. Thus, a first requirement seemed to be fulfilled before knitting haemocompatible tubular conduits from chemically modified PVDF multifilament yarns; unfortunately the available yarns had a too big diameter, which rendered impossible the

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making of enough supple tubular conduits able to be used as vascular substitutes. Except for this application, such yarns seem to not find any other commercial application, which explains the lack of interest from specialised companies to transform PVDF into multifilament yarns.

4.3.3. P(VDF-HFP) case A particular copolymer for which the molar fractions of the two constitutive monomers were 93% and 7%, respectively was available as films prepared by MS Techniques (France) with a 1.78 g/cm3 density. The interest for this copolymer comes from its mechanical properties, which should render it better adapted for making compliant tubings, and from its relatively high radiation resistance. Unfortunately, the availability of multifilament yarns is a bigger concern than it is for PVDF. Nevertheless, the faisability of the modification of this copolymer to endow it with heparin-like properties has been investigated according to the same protocol. During the grafting stage, the styrene can diffuse more easily into this copolymer than it does through PVDF as the latter has a higher level of crystallinity [Fig. 11 clearly shows the presence of 1-mm-large spherulites at the surface of PVDF, whereas such features do not appear at the surface of the P(VDF-HFP)]. In fact, swelling measurements performed with styrene at 60 C on pristine and irradiated polymers, show that the diffusion coefficient is almost 15 times smaller for PVDF than for P(VDF-HFP), this result being attributed to the difference of crystallinity between the two materials [36]. At the same time, the plasticising effect of HFP increases the

Fig. 11. Comparison of PVDF (a) and P (VDF-HFP) (b) surface respective morphologies as shown by scanning electron microscopy.

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mobility of the radicals created upon irradiation and their ability to recombine, which explains a lower global grafting yield of styrene on P(VDF-HFP) in comparison to the grafting yield on PVDF following the same absorbed dose. The previously mentioned authors [36] have investigated in details the role of different parameters: radiation kind (g or SHI) adsorbed dose or fluence for SHI, duration of exposure to the monomer, on the distribution of the grafted polystyrene. The latter may consist in separated islands, when SHI are used at low fluence and the exposure to the monomer is short (Fig. 10), or in a continuous and more or less thick film, which confers to the polymer surface a smooth appearance when g-rays are used and the grafting yield is high (Fig. 12).

5. CONCLUSIONS More than 30 years after it has been proposed for the manufacture of arterial substitutes, microporous PTFE stands more than ever among the biomaterials that are widely used by vascular surgeons. As it may be transformed into tubular conduits, they use the latter to replace arterial segments, or to build femoropopliteal by-pass, or even to build arterio-venous shunts for patients needing recurrent haemodialysis. Its biostability has been established along many years of clinical use, and according to manufacturers, more than 2 millions of microporous Teflon-based vascular prostheses have been implanted all over the world since they have been launched in the market. In the meantime, the original products have suffered very few modifications. Its wall has been made thinner to make its implantation easier, and a gelatin coating may be applied to prevent any bleeding through suture stitches. However, the relative success of such devices is beyond any rational analysis; their compliancy is rather poor and their integration into the host tissues is uneven; moreover, their internal surface may be colonised in vivo but without any positive bioactivity as it may turn thrombogenic again upon biochemical or mechanical stimulation [40]. In fact, despite spectacular clinical results, microporous Teflon-based vascular prostheses do not objectively perform better than the autologous venous graft, which still stands as the gold standard. Clinical studies do not establish as well that these prostheses perform better than polyester-based ones when they are used to build a femoro-popliteal by-pass. If one considers more demanding applications such as the replacement or coronary arteries, the lack of performant prostheses is still pending, and the only chance to fill the gap is to consider other polymeric materials (including other fluorinated polymers) and new strategies aimed at designing vascular substitutes. Among these strategies, those which are based on tissue engineering concepts appears as very promising, and almost ready to be developed. The availability of more compliant fluorinated polymers or copolymers than ePTFE has been evoked in this chapter as well as the availability of methods allowing the functionalisation of their surface to endow the latter with

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Fig. 12. Scanning electron microscopy microphotographies of different polystyrene grafted polyvinylidene difluoride (PVDF) surfaces; the grafting was performed upon gamma irradiation with various absorbed doses; the grafting yield (Y ) was also modulated by the time of exposure of the irradiated surfaces to the styrene. (a) PVDF-gg-PS (Y ¼ 6%), D ¼ 10 kGy; (b) PVDF-gg-PS (Y ¼ 20%), D ¼ 10 kGy; (c) PVDF-gg-PS (Y ¼ 50%), D ¼ 30 kGy.

pro-adhesive properties towards endothelial cells. It is just necessary to develop the appropriate processing techniques, enabling the manufacture of compliant small diameter tubular conduits from these materials, these conduits being able to be sutured by the surgeons to the natural vessels.

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Note from the Editors Partially adapted from ‘‘Biomate´riaux fluore´s pour la chirurgie cardio-vasculaire’’, C. Baquey and R. Guidoin, Actualite´ Chimique, No. 301–302, November 2006. Actualite´ Chimique and Socie´te´ Franc¸aise de Chimie are acknowledged for authorization.