Natural-Based Polyurethane Biomaterials for Medical Applications

Biomaterials are biologically inert or compatible materials placed inside a patient on a longterm or permanent basis. Advances in engineering and a greater availability of synthetic materials triggered the development of engineered polymers for use in biomaterials medical devices. As for other biomaterials, the basic design criteria for polymers used in the body call for compounds that are biocompatible, processable, sterilizable and capable of controlled stability or degradation in response to biological conditions. The interdisciplinary field of biomaterials and tissue engineering has been one of the most dynamic disciplines during the last decades (Durairaj, 2001; Grundke, 2005; Norde, 2006; Ohya, 2002; Pilkey, 2005; Thomson, 2005; Vermette et al., 2001). The selection of the materials used in the construction of prostheses and implants is basically focused on their ability to maintain mechanical, chemical and structural integrity and on various characteristics which allow this function to substitute any organ or tissue properly and exhibit safe, effective performance within the body. Biocompatibility has been defined as the ability of a material to perform with an appropriate host response in a specific application. Any material used satisfactorily in orthopedic surgery may be inappropriate for cardiovascular applications because of its thrombogenic properties. Any deleterious effects may be encountered if used under stress-strain conditions. Biocompatibility of a material can be simulated by comparing its behaviour to reference materials in standardized experimental condition. Biocompatibility is in fact, complex being interpreted as a series of events of interactions happening at the tissue/material interface, allowing the identification of those materials with surface characteristics and/or polymer chemistry more biocompatible; these interactions are influenced by intrinsic characteristics of the material, the confrontational circumstances related to the bioactive and biocooperative responses. Blood contact assays have been developed and include tests investigating the adhesion or activation of blood cells, proteins, and macromolecules such as those found in the complement or coagulation cascade. Other biocompatibility tests have been tentatively proposed and involve analytical testing or observations of physiological phenomena, reactions or surface properties attributable to a specific application such as protein adsorption characteristics. Polyurethanes are one of the most popular groups of biomaterials applied for medical devices. Their segmented block copolymeric character endows them a wide range of versatility in terms of tailoring their physical properties, blood and tissue compatibility.


Introduction
Biomaterials are biologically inert or compatible materials placed inside a patient on a longterm or permanent basis.Advances in engineering and a greater availability of synthetic materials triggered the development of engineered polymers for use in biomaterials medical devices.As for other biomaterials, the basic design criteria for polymers used in the body call for compounds that are biocompatible, processable, sterilizable and capable of controlled stability or degradation in response to biological conditions.The interdisciplinary field of biomaterials and tissue engineering has been one of the most dynamic disciplines during the last decades (Durairaj, 2001;Grundke, 2005;Norde, 2006;Ohya, 2002;Pilkey, 2005;Thomson, 2005;Vermette et al., 2001).The selection of the materials used in the construction of prostheses and implants is basically focused on their ability to maintain mechanical, chemical and structural integrity and on various characteristics which allow this function to substitute any organ or tissue properly and exhibit safe, effective performance within the body.Biocompatibility has been defined as the ability of a material to perform with an appropriate host response in a specific application.Any material used satisfactorily in orthopedic surgery may be inappropriate for cardiovascular applications because of its thrombogenic properties.Any deleterious effects may be encountered if used under stress-strain conditions.Biocompatibility of a material can be simulated by comparing its behaviour to reference materials in standardized experimental condition.Biocompatibility is in fact, complex being interpreted as a series of events of interactions happening at the tissue/material interface, allowing the identification of those materials with surface characteristics and/or polymer chemistry more biocompatible; these interactions are influenced by intrinsic characteristics of the material, the confrontational circumstances related to the bioactive and biocooperative responses.Blood contact assays have been developed and include tests investigating the adhesion or activation of blood cells, proteins, and macromolecules such as those found in the complement or coagulation cascade.Other biocompatibility tests have been tentatively proposed and involve analytical testing or observations of physiological phenomena, reactions or surface properties attributable to a specific application such as protein adsorption characteristics.Polyurethanes are one of the most popular groups of biomaterials applied for medical devices.Their segmented block copolymeric character endows them a wide range of versatility in terms of tailoring their physical properties, blood and tissue compatibility.lead wire insulations, intra-aortic balloons, catheters and artificial hearts, because of their generally favorable surface physical properties, together with their fairly good biocompatibility and haemocompatibility characteristics, (Lamba et al., 1997;Lelah & Cooper, 1986;Plank et al., 1987).The balance between the surface hydrophilic and hydrophobic properties is important for achieving an enhanced biocompatibility of polyurethanes.Plasma treatments or other types of stimuli may alter the surface energy of most polymers, thus changing their surface polarity, hydrophilicity and adhesiveness, (Desai et al., 2000;Ozdemir et al., 2002;Ramis et al., 2001).In Table 1  The measurement methods used for determination of surface tension are based on contact angles between the liquid meniscus and the polyurethane surface.The contact angle is a measure of the ability of a liquid to spread on a surface.The contact angle is linked to the surface energy and so one can calculate the surface energy and discriminate between polar and apolar interactions.For the calculation of the surface tension parameters, the geometric mean method (Eqns.( 1) and ( 2)), (Kälble, 1969;Owens & Wendt, 1969;Rabel, 1977) the acid/base method (LW/AB) (Eqns.( 3)-( 5)), (Van Oss et al., 1988a;Van Oss et al., 1988b;Van Oss, 1994), and theoretical method based on the structure-property relationship considering the group contribution techniques (Eqn.( 6)), (Bicerano, 1996) , were used.
where γ is the total surface tension, E coh the cohesive energy and V the molar volume.
According to the geometric mean method, the solid surface tension components were evaluated with Eqn. ( 1), (Van Oss et al., 1989) using the known surface tension components, (Erbil, 1997;Rankl et al., 2003;Strom et al., 1987) of different liquids from Table 5 shows the contribution of the polar component to the total surface tension obtained from the geometric mean method GM for untreated and plasma treated polyurethanes.The total surface tension was estimated from the structure-property relationship, according to Eqn. (6) in the following steps, (Bicerano, 1996): 1. Calculation of the zeroth-order connectivity indices o χ and o χ v and of the first-order connectivity indices 1 χ and 1 χ v , according the values of the atomic simple connectivity indices and of the valence connectivity indices (Table 6).2. Calculation of cohesive energy, by two methods, by applying the group contributions of Fedors, (Bicerano, 1996;Van Krevelen, 1990) and those of Van Krevelen and Hoftyzer, (Bicerano, 1996;Van Krevelen, 1990), (Table 7).3. Calculation of the molar volume at room temperature (298 K), (  and γ (2) , from the theoretical data calculated for cohesive energies, E coh( 1) and E coh(2) , and molar volume, V, for studied polyurethanes The theoretical results are closed to the experimental values, derived from the contact angle measurements.
The hydrophobe-hydrophile balance of untreated and plasma treated polyurethanes has been evaluated by calculation of free energy of hydration, ΔG w .The ΔG w values were obtained from Eqn.( 7), (Faibish, 2002): where γ lv is the total surface tension of water from Generally, the literature (Faibish, 2002;Van Oss, 1994) of the field mentions that for ΔG w < -113 mJm -2 the polymer can be considered more hydrophilic while when ΔG w > -113 mJm -2 it should be considered more hydrophobic.High frequency cold plasma treatment modifies ΔG w indicating that the surface becomes more hydrophilic in the case of PPG-HPC sample and less hydrophilic in the case of PEA-HPC and PTHF-HPC samples.Solid-liquid interfacial tension is defined with the following relation: Free energy of hydration and interfacial tension are very important in that they determines the interactional force between two different media and controls the different processes: stability of the colloidal aqueous suspensions, dynamic of the molecular self-assembling, wetability of the surface, space distribution and adhesiveness.The biological and chemical processes, which take place at the level of the surface of the implant, depend on the interfacial interactions between solid and liquid (water).
(1)When the blood-biomaterial interfacial tension is high, the blood proteins will be anchored on many points on the surface, they strongly interact with the surface and thus the solid-liquid interfacial tension decreases.Consequently, the proteins change their conformation.A new interface is formed, between the protein surface and sanguine plasma.
(2)When the blood-biomaterial tension is relatively low, the force, which determines the protein adsorption, will be smaller.Conformation of the proteins initially adsorbed is similar to that found for the proteins in solution.Therefore, the interfacial tension between the protein surface and the sanguine plasma will not be high, not being an appreciable force able to determine the adsorption of sanguine components.This corresponds to a better compatibility of the biomaterial surface with blood comparing with the case (1).The surface of the biomaterial must reduce to minimum the blood-biomaterial interfacial tension such as the modification of the initially adsorbed proteins to be little.Although, apparently an interfacial tension equal to zero would be ideal for realization of the blood compatibility, however this is not desirable in view of the mechanical stability of the blood-biomaterial interface.It is generally considered that the blood-biomaterial interfacial tension should be 1-3 mN/m for a good blood -biomaterial compatibility, as well as a good mechanical stability of the interface.
The values for solid-liquid interfacial tensions are given in Table 8 for untreated and plasma treated samples.It can be observed that the interfacial tensions are in general low, and after plasma treatment become even lower.Moreover, 1 mN/m < γ sl for PTHF-HPC and PPG-HPC samples treated in plasma < 3 mN/m which is required for a good biomaterial.An important goal in material science, biochemistry, cell biology and bioanalytics is to establish a relationship between surface morphology and surface properties of a polyurethane composite on one hand and on the other hand the interconnective interactions synthetic polymer matrix-extracellular matrix natural polymer components (Macocinschi et al., 2010a).The polyurethane was modified with natural polymers aiming the improvement of the surface and bulk properties, to confer good biomaterial qualities.The added natural extracellular matrix polymers determine an increase of the surface tension value for all the biocomposite samples in comparison with starting polyurethane.The fact that the total surface tension values experimentally evaluated for starting polyurethane (untreated and treated in plasma) are less than the theoretical one calculated on the basis of quantitative structure-properties relationships (Bicerano), which can be explained by the complex morphology of polyurethanes with microdomain segregation and network.The complex network morphology gives stability against plasma action.For all the tested biocomposites the values of the total surface tension after plasma treatment are higher, therefore the samples become more hydrophilic except one sample containing chondroitin sulfate, which suggests that chondroitin sulfate component induces further crosslinks and network formation and hence a reducing of hydrophilicity.The evaluated surface tension parameters show that the biocomposites have biomaterial qualities through their increased hydrophilicity.Moreover, plasma treatment decreases the hydration energy while the interfacial tensions fall within 1-3 mN/m, interval required for a good biomaterial.Biocompatible interfaces were constructed based on non-toxic, hydrophilic assembled macromolecular networks, as viable platforms for the affinity of biomolecules.The mechanical strength of the biocomposites decreases in comparison with starting polyurethane but it is higher than that of carotid porcine arteries.

Scanning electron microscopy
In

Water sorption
The biocompatibility of the materials depend on their ability to swell in aqueous media.A high water level on the surface of the biomaterial provides a low interfacial tension with blood, thus reducing fibrinogen adsorption, cell adhesion and clot formation, (Abraham, 2002;Faibish, 2002;Van Krevelen, 1990;Wang, 2004).The results are presented in Fig 2 .It is observed that the water uptake is given by the PEA-PU sample (reference polyurethane sample without hydroxypropylcellulose, PEA/MDI/EG, M n =109.613,M w /M n =1.3), PEA-HPC and PTHF-HPC samples (151 %, 140 % and 167 %, respectively ) and in a less extent by the PPG-HPC (92 %) due to its less polar soft segment having the lateral -CH 3 substituent which confer a different geometry to the polyurethane internal microporous structure, unfavorable for water uptake.

Fibrinogen adsorption
The  It is observed that the amount of fibrinogen adsorbed from sanguine plasma is less in comparison with that in solution, for all the materials except PPG-HPC, for which the differences are not significant.Also, it is remarked that prothrombin time stand in physiological normal limits (Table 9) so studied polyurethane samples did not affect the clot formation mechanisms.The significant differences between adsorptions of the fibrinogen solution and sanguine plasma, suggest that the fibrinogen adsorption properties of the polyurethane samples, under physiological condition are affected by the concurrent affinities for other plasma proteins, which do not disturb the haemostatic mechanisms.Probably among the plasma proteins that can concur with fibrinogen is albumin, which was investigated in (Lupu et al., 2007a), and it was found that the adsorption value of a sub-physiological solution of serum albumin, (3 mg/ml) to PEA-PU materials was 0.  10 it is observed that at first immersion θ adv are less than 90 o which concludes that the polyurethane materials show hydrophilicity.These film samples have been prepared through precipitating in water making them microporous and enough hydrophile at surface through orientation of the hydrophile groups towards surface.The evaluated receding contact angles are 44-48 % less, except PEA-PU which is the polyurethane without HPC in its composition.Thus, HPC reduces the θ rec.At the second immersion the hysteresis for PEA-PU is significantly less comparing with all samples having HPC.It appears that a polar material induces a low hysteresis, and the polarity of the soft segment in the polyurethane influences the results, considering additionally at second immersion the water up-take within the microporous layer disposed at surface.Advancing contact angles at second immersion are less than at first immersion, thus the material is better wet.Of all the polyurethane samples, PTHF-HPC evidences the lowest dynamic contact angles, both advancing and receding angles, at first and second immersion, which recommends it for biomedical applications.
In Table 11  When PEA-HPC is solved, due to cellulose derivative, OH groups and hydrogen bonding, the decrease is found to be somewhat higher.For polyether-urethanes, PTHF-HPC and PPG-HPC, there is found a much more significant Δ γ (21.69 %, respectively 20.96 %), and this is due to the polyether nature of the soft segment.This can be explained by the fact that when polar polymers like poly(ester urethane)s are solved in a polar solvent like DMF reduces in some extent the surface tension, while a somewhat less polar polymer like poly(ether urethane)s in the same solvent reduces more.

Dynamic mechanical thermal analysis (DMTA)
Polyurethanes are characterized by their high and heterogeneous morphology and by important intermolecular forces that determine their mechanical behavior.In such way that the forces applied to polymers in general and the deformations produced by those are not thoroughly local.Consequently the response of the polymer to the foreign solicitations is extended in a wide time interval (relaxation time), originating its peculiar viscoelastic behavior.While energy supplied to a perfectly elastic material is stored and a purely liquid dissipates it, polymeric materials dissipate a part of energy that excites to them.DMTA is a sensitive technique used to study and characterize macroscopic responses of the materials as well local internal motions.By monitoring property changes with respect to the temperature and/or frequency of oscillation, the mechanical dynamic response of the material is referred to two distinct parts: an elastic part (E', storage modulus) and a viscous component (E", loss modulus).The damping is called tan delta, loss factor or loss tangent, tan δ =E"/E'.With increasing temperature different physical states are revealed: glass, leathery, rubbery and elastic of rubbery flow, viscous flow.The glass transition is easily identified from dynamical mechanical data because of the sharp drop in storage modulus, peaks of loss dispersion modulus or tan δ.Tan δ peak may occur at higher temperatures than those given by E' drop or E" peak, because it responds to the volume fraction of the relaxing phase, its shape and height depends on the amorphous phase, being a good measure of the 'leather like' midpoint between glassy and rubbery state, (Sirear, 1997).The glass transition temperature (Tg) is often measured by DSC (Differential Scanning Calorimetry), but the DMTA technique is more sensitive and yields more easily interpreted data.This is usual as the degree of dependence is specific to the transition type.DMTA can also resolve sub-Tg transitions, like beta, gamma, and delta transitions, in many materials that the DSC technique is not sensitive enough to pick up.The magnitude of the low temperature relaxations is much smaller than that of α-relaxation considered as the glass transition.These relaxations are due to local mode (main chain) relaxations of polymer chains and rotations of terminal groups or side chains, or crankshaft motion of a few segments of the main chain.
In literature for polyurethanes based on poly(epsilon caprolactone) and 1,4-butane diisocyanate with different soft segment lengths and constant hard segment length it was evidenced by DMTA additional transitions at room temperature due to crystalline fraction of PCL while the hard segment crystallinity influence the rubber plateau, (Heijkants et al., 2005).For poly(ether urethane) networks prepared from renewable resources: epoxidized methyl-oleate polyether polyol and 1,3-propandiol by using L-lysine diisocyanate, it was found very well differentiated both from DSC and DMTA, two glass transitions for the soft segment (-17-1 o C; 9-22 o C) and respectively hard segment (35, 44 o C; 45, 58 o C) explainable through phase segregated morphology, (Lligadas et al., 2007).Three kinds of polyurethane mixed blocks (polycaprolactone glycol, polypropylene glycol, polytetramethylene glycol) and 4,4'-diphenylmethane diisocyanate extended with 1,4-butane diol were studied by DMTA and it was revealed that soft chain mobility affects the glassy state modulus.From E' and tan δ graphs the T g values are less than -50 o C, (Mondal, 2006).Gao and Zhang, (Gao & Zang, 2001) found IPNs as a novel kind of material with complex internal friction behavior and thermal dynamic incompatibility: for their semiinterpenetrating networks of castor oil polyurethane and nitrokonjac glucomannan, glass transition temperatures (T g ≤ 50 o C ) increased with molecular weight of the latter component which affects the storage modulus and shape and position of tan δ by changing the degree of order and motion of molecules through concentration fluctuations of molecular structural units.For graft-interpenetrating polyurethane networks and natural polymers such as nitrolignin, (Huang & Zhang, 2002), the influence of the NCO/OH molar ratio was studied by DSC and DMTA revealing T g = -4.09-23.97o C (DSC) and respectively (T g = 6.3-31.1 o C) which increase with NCO/OH molar ratio through formation of three-dimensional allophanate or biuret networks.DMTA and DSC investigations on new blends of hydroxypropylcellulose and polyurethane (poly (ethylene glycol) adipate -4,4'-diphenylmethane diisocyanate-ethylene glycol) reveals glass transitions ranging between (-17-11.8 o C) by DMTA and T m for soft segments (41.6-49.7)o C by DSC, (Raschip et al., 2009).Characteristic temperatures for the studied samples determined from DMTA curves (Figs. 4,5,6) are listed in Table 12.
Storage modulus E' is a measure of the stiffness of the material, (Zlatanic et al., 2002).It appears from the Fig. 4 that the following order of E' values at its drop (at the glass transition from glassy to leathery state) can be stated: E' PEA-HPC > E' PPG-HPC >E ' PEA-PU >E' PTHF- HPC .This is related to the crystalline domains or physical/chemical network/entanglements which constraint molecular motions in amorphous state.For all the samples E' is less than 10 9 Pa, being generally known that when E' > 10 9 Pa the material is glassy.In addition we recall that for an amorphous linear polymer the decline of E' in the glass transition range amounts three orders of magnitude in a narrow temperature span.In particular for our samples the decline found for E' is about two orders of magnitude, polyester urethanes are expected to be stiffer, than polyether, cellulose derivative induces also stiffness to the material, while lateral methyl groups in amorphous atactic PPG provokes constraints in the mobility of the soft segment.PTHF macromolecular chain is more mobile with a low stiffness and low tan δ.The glass transition of the soft segment (SS) was determined from DMTA curves as follows: from the intersection of tangents to the E' (log E') curves from the glassy region and the transition "leathery" region, from E" (log E") peaks and tan δ peaks.T g of the soft segment from DSC was evaluated from the second heating scan.It can be noticed that T g from DSC are close to those from DMTA, for poly(ester urethanes) while for poly(ether urethanes) are different and this can be explained by the sensitivity of DMTA technique to the mobility of the macromolecular segment.And we referred here to the T g values evaluated from E' (log E') graphs.The log E' or log E" vs. E' or E" evidence better the biphasic behaviour of the samples by revealing the melting phenomena related to the soft and hard segment.We remark slope changes on log E' descent and the right edge of the E" peak has a descent trend which indicate a possible overlapping of the melting of the soft phase with a glass transition of the hard phase.The broadening of the glass transition reveals large distribution of the relaxation times that implies a heterogeneous structure with a soft phase constraint by a hard phase which reduces its mobility.Poly(ether urethanes) samples evidence secondary relaxations of the soft segment, attributed to local relaxations in glassy state, which may imply few methylene groups or the motion of -CH 3 attached to the backbone or crankshaft motion: for PTHF-HPC sample below glass transition of the soft segment β relaxation is evidenced by tan δ graph (T β = -51.Tm(SS) -melting point of soft segment; Tm(HS)dec.-melting point of hard segment accompanied by decomposition; T β -secondary transitions below glass transition considered as α-transition (T α ).
Table 12.Characteristic temperatures for the studied samples determined from DMTA and DSC curves From Fig. 6 tan δ values at glass transition peaks show that E' is almost 5E" for all the samples, except PTHF-HPC for which E' is almost 10 E".This result evidences that for all the samples the elastic modulus component is more important than the viscous modulus one.

Mechanical properties
Biological materials have a wide range of mechanical properties matching their biological function.This is achieved via assembly of different size building block segments (soft and hard) spanning many length scales.Due to specific chemical versatility of polyurethanes, different morphologies at different length scales can be obtained and thus different physical properties which satisfy diverse clinical needs have been achieved.The modulability of mechanical properties make polyurethanes excellent candidates for applications in soft tissue engineering.Because of the strong tendency of rigid aromatic moieties to pack efficiently and the presence of hydrogen bonding between urethane and urea groups they tend to selforganize to form semi-crystalline phases within the polymer macromolecular assembly.As the elasticity of the polymers depends on their degree of crystallinity and the degree of hard segment segregation, it is clear that the selection of the diisocyanate monomer will be one of the key parameters that influence polyurethane mechanical characteristics.
The resulted tensile properties are tabulated in  The stress-strain curves of the studied polyurethanes are plotted in Fig. 7 for dry film samples and in Fig. 8 the stress-strain curves of the film samples previously conditioned in warm (37 o C) saline water (NaCl, 0.9 % w/v, pH=7.4) for 24 h, then blotted with absorbent filter paper, are presented.We compare in this way, the mechanical properties of the films in dry state vs. physiological condition.The influence of the physical and chemical cross-links on the elastic behaviour of the polyurethanes is investigated by using Mooney-Rivlin equation, ( 9), for rubbers, (Sekkar et al., 2000;Spathis, 29) : where σ is the stress, λ is the extension ratio (L/L 0 ) and C 1 , C 2 are the Mooney-Rivlin constants.
From the stress-strain experimental data, the Mooney-Rivlin curves are plotted (Fig. 9 and Fig. 10) and the values of C 1 and C 2 are obtained (see Table 13).C 1 can be obtained by extrapolating the linear portion of the curve to λ -1 = 0, and C 2 from the slope of the linear portion.The elastic behaviour depends on the size and the distribution of the hard domains into the soft matrix and is more or less reflected in the deviation from the Mooney-Rivlin equation.Therefore, C 1 is different for HPC polyether-and polyester-urethanes (C 1 is lower for polyester-than polyether-urethanes).At higher strains a disrupting of the physical crosslinks is taking place and C 1 becomes lower.C 2 is not so sensitive relative to the nature of soft segment but shows the presence of both physical and chemical cross-links specific to a polyurethane network.The effect of conditioning in saline water (0.9 % w/v, 24 h, 37 o C) is clearly revealed by C 1 which is found lower for all the samples evidencing that after conditioning the physical network is affected as well as its flexibility.
The toughness representing the energy absorbed before the sample breaks is higher as expected for PEA-HPC sample than for the PPG-HPC and PTHF-HPC samples, both for dry and conditioned samples.Moreover it can be noticed that the cellulose derivative makes that this absorbed energy to be much higher in case of dry samples.Hydration of samples leads to the increase of toughness except PPG-HPC sample due to amorphous and atactic soft segment structure.Hydration of the semicrystalline, more or less ordered structures like polyurethanes, have as main result the disrupting of the physical bonding, and the plasticizing of the biopolyurethane matrix, affecting strain behaviour.Water may penetrate within interstitials of the microporous structure favoring biological interactions.The heat is also important in softening the material acting upon the soft segment and physical hydrogen bonding.From Table 13 one can notice that for polyether-urethanes (PTHF-HPC, PPG-HPC), Young modulus decreases after conditioning in saline water at 37 o C for 24 h, the material achieve more flexibility, which may bring as advantage for realization of biopolyurethane tissular structures, such heart valves and leather grafts.Hydrogen bonding is known as an important driving force for the phase separation of a hard segment from a soft-segment matrix consisting of polyether or polyester polyol.The separated hard segment acts as physical cross-links and filler particles for the soft-segment matrix.Microphase separation of segmented polyurethanes is probably the single most influential characteristic of these materials.The degree of phase separation plays a key role in determining mechanical properties and blood compatibility, (Yoon et al., 2005).The heterogeneous morphology of polyurethanes will determine the surface composition exposed to a polar environment (water or blood) or to a nonpolar environment (air or vacuum).The surface segregation phenomenon reflects the difference in surface energy between the polar and nonpolar components, (Lupu et al., 2007b).The surface composition constitutes a crucial parameter for a biomedical material in contact with blood.The mobility of polymer chains coupled with environmental changes can lead to surface composition and properties that are time-dependent and dependent on the contacting medium, temperature, that the polymer experiences.It is, however, an unresolved question as to whether an air-stored polyurethane surface indeed adapts to the aqueous environment on biomedical usage.As the polyurethane biomaterial is placed into contact with a physiological medium, such as blood or tissue, its surface layers will undergo motions in order to accommodate the new interfacial situation.In contact with aqueous environments, it is obviously favorable for hydrophilic constituents of the polymer to become enriched at the interface, (Vermette & Griesser, 2001).
For crosslinked blends of Pellethene and multiblock polyurethanes containing phospholipids, (Yoo & Kim, 2005), it was found for elastic modulus values ranged between 21-47 MPa, whereas for biomaterials based on cross-linked blends of Pellethene and multiblock polyurethanes containing poly(ethylene oxide) it was found for elastic modulus values ranged between 85-246 MPa, (Yoo & Kim, 2004).

Platelet adhesion
In vitro platelet adhesion experiments were conducted to evaluate the preliminary blood compatibility.It is very well known that the surface properties particularly platelet adhesion of the biomaterials is very important with respect to their haemocompatibility, especially when they are used as cardiovascular devices, (Park et al., 1998).
It is well known that when blood is in contact with a synthetic material, firstly the latter one adsorbs onto its surface blood plasma proteins, and secondly attract and activate the thrombocytes.In function of the type of the adsorbed plasma proteins (fibrinogen or albumins) this phenomenon can exert in a more or less extent.In case of the preferentially fibrinogen adsorption (important protein in endogenous haemostasis), platelet adhesion is increased followed by thrombus activation and clot formation.In case of albumin absorbtion platelet adhesion is diminished which confer to the surface a thromboresistant character, (Bajpai, 2005;Wang et al., 2004).In Fig. 11 platelet adhesion corresponding to the studied polyurethane samples is given.It can be noticed that when comparing PEA-PU with PEA-HPC the no. of adhered thrombocytes is much lower for PEA-HPC and this is due to the cellulose derivative which confers improved biomaterial qualities.Among the PPG-HPC, PTHF-HPC and PEA-HPC samples the most promising thromboresistant ones are PTHF-HPC and PEA-HPC.In our previous paper, the fibrinogen adsorption tests, (Macocinschi et al., 2009a), revealed that PTHF-HPC and PEA-HPC polyurethanes proved to be indicated for thromboresistant devices, and the polyether-urethane PTHF-HPC proved to have more relevant haemocompatible material qualities.

Conclusion
The effect of the chemical structure of some cellulose derivative based segmented biopolyurethanes and polyurethane biocomposites containing extracellular matrix components on surface properties (static and dynamic contact angle measurements), on dynamical mechanical thermal analysis, mechanical properties was investigated.Specific biological tests were performed (platelet adhesion, protein adsorption tests of fibrinogen) and the results obtained recommend them as biomaterials with enhanced biocompatibility for biomedical applications.Thus, the proper understanding of blood compatibility and the development of prosthetic materials require the investigations of both natural and synthetic materials.Extensive in vitro, ex vivo, and in vivo testing in suitable animals and related differences between the properties of blood components of animals and humans are essential.Future research should be directed in order to clarify the influence of anatomical site of implantation on the behaviour of the implant, with a goal towards the development of site specific implants.

References
Fig 1 are illustrated the SEM micrographs corresponding to the treated and untreated polyurethane samples.It is obvious that plasma caused a change in the surface morphology and etching effects are observed.

Fig. 1 .
Fig. 1.SEM images of the polyurethane samples untreated and treated in HF cold plasma (PEA-HPC (a,b); PTHF-HPC (c,d); PPG-HPC (e,f)) experimental data related to the amount o f a d s o r b e d f i b r i n o g e n b e f o r e a n d a f t e r incubating of polymers with a physiological solution of fibrinogen (3.00 mg/mL) and sanguine plasma (2.98 mg/mL) are presented in Fig 3.

Fig. 2 .
Fig. 2. Weight of polymer sample in dry and maximum hydrated state 4 o C) while for PPG-HPC sample β and γ relaxations arised probably at the level of the main chain and lateral -CH 3 groups (E" graph T γ = -128 o C and T β =-81.1 o C; log E" graph T γ = -131.6o C and T β =-82.5 o C; tan δ graph T γ = -129.1 o C and T β = -84.4

Fig. 4 .
Fig. 4. Storage modulus as a function of temperature for the studied samples

Table 1 .
are provided the compositional parameters and the average molecular weights values, polydispersity indices (GPC).Compositional parameters, number-average molecular weights, polydispersity indices

Table 2 .
Contact angle degrees of different liquids and polyurethane samples before and after plasma treatment Table 2 lists the contact angles between double distilled water, ethylene glycol, or CH 2 I 2 and polyurethane samples, before and after effecting of high frequency cold plasma treatment.

Table 3
Table3and the contact angles from Table2.The total surface tension was calculated with Eqn.(2)..Surface tension parameters (mN/m) of the liquids used for contact angle measurements Table4shows the surface tension parameters for both untreated and plasma-treated polyurethane samples, according to the geometric mean method and to the acid/base method.In this table it was considered that LW sv γ is equivalent to d sv γ of the geometric mean method, the mean values of sv γ − and sv γ + were calculated with Eqn.(3).Also, the total surface tension was calculated with Eqn.(2).Following the plasma treatment the disperse component of surface tension, d sv γ , increases in absolute value, while the polar component surface tension p

Table 8
Table 3 and θ water is contact angle of water with polyurethanes.The results are presented in Table 8.
. Surface free energy between polyurethane and water and interfacial tensions for untreated and plasma treated samples

Table 9 .
3±0.06 mg/cm 2 .Prothrombin time before and after the contact sanguine plasma with polyurethane samples2.5DynamiccontactangleInTable10are presented the dynamic contact angle values (advancing θ adv, o receding θ rec, o and hysteresis H, %) of the film samples in contact with water.

Table 10 .
Dynamic contact angles values of the film samples in contact with waterFrom Table

Table 11 .
are given the experimental values of the surface tension for the polyurethane solutions in DMF and DMF as solvent, by using Wilhelmy plate method.Solutions of the polyurethane sample 4g/dL in DMF were employed and DMF as a starting solvent.The experimental values of the surface tension for the polyurethane solutions in DMF and DMF as solvent It is obvious in Table11that γ DMF decreases only with 7.62 % by solving the PEA-PU polymer.