**3.3.** *In vitro* **and** *in vivo* **performances of PU/HPC membranes**

The biocompatibility of PUs are widely discussed and questioned, mostly in the past. In the last two decades new generation of PUs that combine mechanical advantages with the biological performances emerged (Gisselfa et. al., 2002; Jordan & Chaikof et. al., 2007; Jun et. al., 2005; Kavlock et. al., 2007; Parveen et. al., 2008). For many years it has been considered that PUs biocompatibility is spotless due to their products of degradation, e.g., aromatic polyamines. As it is well known for the most part of biocompatible materials, the life time of their *in vitro* functionality is quite short. This is a consequence of their intrinsic physicochemical properties, on one hand, and of the tissue action on the material, on the other hand (Anderson, 2001; Guelcher, 2008; Shen & Horbett, 2001).

#### *3.3.1. Oxidative in vitro behavior*

Oxidative degradation of PUs caused by hydrolytic or enzymatic mechanism was intensively discussed (Christenson et. al., 2004; Guelcher, 2008; Gary & Howard, 2002; Sutherland et. al., 1993). First of all, PUs designed for tissue-contact devices undergo hydrolytic degradation as a result of watering with physiological solutions. This process has an impact especially on poly(ester-urethane)s that can generate hydroxy-acids, being susceptible to induce reactive oxygen species (ROS) production following the materialtissues interaction. By means of this mechanism, PUs can be implied in the sustained oxidative degradation and a wide range of pathological states.

As it is well known, ROS can trigger subtle mechanisms responsible for diseases generation through the peroxidation of cell membrane lipids and DNA damage (Marnett, 2002; Tribble et. al., 1987; Yagi, 1987). The most susceptible organs to oxidative aggression are the heart, vessels, lung, gut, liver, brain and nerves (Ames et. al., 1993; Förstermann, 2008; Paradis et. al., 1997; Rahman et. al., 2002; Sayre et. al., 1997).

In a normal body state, ROS appear constantly as a result of some biological errors or as a consequence of some short living reactive intermediate products generated by the cell aerobic metabolism. Endogenous enzymatic and nonenzymatic pathways are responsible for the formation of free radicals. These pathways are balanced by two endogenous antioxidant pathways, which form the TAS (see fig. 7).

While some harmful material characteristics can be marked as cytotoxic or proinflammatory by standard testing, others, such as oxidative stress (that causes long-time material failure), are undetectable by using short period testing. Thus, well known biocompatible materials were found to display surface alteration or cracking after long-time implantation. Adding antioxidant compounds to materials can improve their resistance against tissue degradation (Oral et. al., 2006; Stachelek et. al., 2006; Wattamwar et. al., 2010).

212 Polyurethane

biomedical applications.

structures and improve their resistance for autoclaving procedures as important step in ready to use biomaterials preparation. From all the data presented in this section, one can say that the more hydrophilic PU-PTHF/HPC membrane could be the most appropriate for

The biocompatibility of PUs are widely discussed and questioned, mostly in the past. In the last two decades new generation of PUs that combine mechanical advantages with the biological performances emerged (Gisselfa et. al., 2002; Jordan & Chaikof et. al., 2007; Jun et. al., 2005; Kavlock et. al., 2007; Parveen et. al., 2008). For many years it has been considered that PUs biocompatibility is spotless due to their products of degradation, e.g., aromatic polyamines. As it is well known for the most part of biocompatible materials, the life time of their *in vitro* functionality is quite short. This is a consequence of their intrinsic physicochemical properties, on one hand, and of the tissue action on the material, on the other hand

Oxidative degradation of PUs caused by hydrolytic or enzymatic mechanism was intensively discussed (Christenson et. al., 2004; Guelcher, 2008; Gary & Howard, 2002; Sutherland et. al., 1993). First of all, PUs designed for tissue-contact devices undergo hydrolytic degradation as a result of watering with physiological solutions. This process has an impact especially on poly(ester-urethane)s that can generate hydroxy-acids, being susceptible to induce reactive oxygen species (ROS) production following the materialtissues interaction. By means of this mechanism, PUs can be implied in the sustained

As it is well known, ROS can trigger subtle mechanisms responsible for diseases generation through the peroxidation of cell membrane lipids and DNA damage (Marnett, 2002; Tribble et. al., 1987; Yagi, 1987). The most susceptible organs to oxidative aggression are the heart, vessels, lung, gut, liver, brain and nerves (Ames et. al., 1993; Förstermann, 2008; Paradis et.

In a normal body state, ROS appear constantly as a result of some biological errors or as a consequence of some short living reactive intermediate products generated by the cell aerobic metabolism. Endogenous enzymatic and nonenzymatic pathways are responsible for the formation of free radicals. These pathways are balanced by two endogenous

While some harmful material characteristics can be marked as cytotoxic or proinflammatory by standard testing, others, such as oxidative stress (that causes long-time material failure), are undetectable by using short period testing. Thus, well known biocompatible materials were found to display surface alteration or cracking after long-time implantation. Adding

**3.3.** *In vitro* **and** *in vivo* **performances of PU/HPC membranes** 

(Anderson, 2001; Guelcher, 2008; Shen & Horbett, 2001).

oxidative degradation and a wide range of pathological states.

al., 1997; Rahman et. al., 2002; Sayre et. al., 1997).

antioxidant pathways, which form the TAS (see fig. 7).

*3.3.1. Oxidative in vitro behavior* 

**Figure 7.** Schematic representation of the oxidative/antioxidative balance with enzymatic and nonenzymatic tissue pathways

Antioxidant defensive systems are present in both cells and extracellular environment. SA molecules are the most important antioxidants in blood. Due to their high concentration and polyvalent possibilities to fit with oxygen free radicals, SA molecules are considered to be the main plasmatic components of defence that assure neutralisation of more than 70% of ROS (Bourdon & Blache, 2001).

Assigning to SA molecules the main role in protective effect, we analysed the interaction of PU/HPC membranes with blood plasma, following the plasma antioxidant status. To define the importance of SA adsorption on material surface, the membranes were incubated at 37 oC in blood plasma and TAS was measured periodically. The results are shown in Fig. 8.

Two PU samples (PU-PEGA and the more hydrophobic PU-PPG/HPC) had significant tendency to quickly decrease TAS activity in the first 48 hours. Due to the complexity of TAS, it is difficult to speculate on the mechanism by which the decreasing phenomenon arises and certainly more examinations are needed. However, one can suppose that PU-PEGA alter the TAS activity as a result of plasma pH modification that leads to sustained free radical generation in the presence of the material. The mechanisms by which TAS activity is lowered after PU-PPG/HPC incubation could not be related directly to SA antioxidant activity, but to some other oxidant pathways that need further investigations.

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of the Improved Polyurethane Membranes for Medical Applications 215

A negative charge acts as an activator of plasmatic factor XII (Hageman factor) that involves contact system and intrinsic coagulation pathway (Zhuo et.al. 2006). This mechanism also involves high molecular weight and positively charged kininogen (HMW) and plasma thromboplastin (factor XI). Contact mechanism is tightly related to inflammatory events because some intrinsic pathways factors are direct activators of neutrophils. Whole mechanism of contact blood coagulation is still unclear. Some authors hypothesized that it can also be induced by adsorbed FB (Colman & Schmaier, 1997) and hydrophobic surfaces (Zhuo et. al., 2006). As for intrinsic coagulation mechanism of thrombus formation, this can also be activated by negatively charged low density lipoproteins (LDL), the molecules that adhere to the vessel walls in some pathologic conditions associated with cardiovascular risks (Krieter et. al., 2005). This possible mechanism should be taken into account because almost all pathological situations in which blood-assisted devices are used are accompanied by high level of cardiovascular risk factors (high level of LDL, cholesterol, triglycerides and

**Figure 9.** Contact coagulation and extrinsic (tissue factor) coagulation blood pathways

The obtained results are summarised in Table 2.

PUs are promising materials for implantable and non-implantable blood-interacting devices. They combine an increased elasticity with good mechanical resistance. For haemocompatibility evaluation of PU/HPC materials discussed above, the haemolytic and thrombotic potentials were determined by standard and adapted methods (see section 2).

modified blood pressure).

**Figure 8.** TAS evolution after PU/HPC incubation in blood plasma at 37 oC

#### *3.3.2. In vitro haemocompatibility*

Haemocompatibility involves compatibility with blood cells and blood plasma in other words, nonhaemolytic and nonthrombogenic behavior. Haemolysis is a mechanism by which erythrocytes (red blood cells) are destroyed through cell membrane lyses. Erythrocyte membrane lyses may occur as a result of environment pH modification or by cytotoxic action on erythrocyte membrane. Thus, both lipid (by lipid peroxidation) and protein (by protein modification) compounds can be affected.

Thrombogenesis is a complex phenomenon by which thrombus is formed by blood clotting. As a physiological event, haemostasis implies the activation of the enzymatic cascades in which three main factors are involved – vascular, cellular and plasmatic (Edmunds, 1998).

A synthetic material can induce haemostasis activation by surface charge, hydrophobicity and/or released products of degradation. It is widely recognised that both positively and negatively charged as well as hydrophobic surfaces can induce thrombus formation. This can be explained by involvement of several mechanisms as presented in Fig. 9.

A positive charge can be favourable for FB adsorption, followed by its conformational modification and adhesion of platelets and leukocytes (monocytes). Adherent cells are activated and they release numerous molecules that lead finally to FB cleavage with fibrin network formation (clot). Among platelet secreted factors are platelet thromboplastin, fibrin stabilizing factor, serotonin, anti-heparin factor, and others. Adherent (activated) monocyte releases thrombogen tissue factor (TF). Mechanism triggered by positive and hydrophobic surfaces is mainly related to extrinsic coagulation pathway (B. Furie & B. C. Furie, 2008).

A negative charge acts as an activator of plasmatic factor XII (Hageman factor) that involves contact system and intrinsic coagulation pathway (Zhuo et.al. 2006). This mechanism also involves high molecular weight and positively charged kininogen (HMW) and plasma thromboplastin (factor XI). Contact mechanism is tightly related to inflammatory events because some intrinsic pathways factors are direct activators of neutrophils. Whole mechanism of contact blood coagulation is still unclear. Some authors hypothesized that it can also be induced by adsorbed FB (Colman & Schmaier, 1997) and hydrophobic surfaces (Zhuo et. al., 2006). As for intrinsic coagulation mechanism of thrombus formation, this can also be activated by negatively charged low density lipoproteins (LDL), the molecules that adhere to the vessel walls in some pathologic conditions associated with cardiovascular risks (Krieter et. al., 2005). This possible mechanism should be taken into account because almost all pathological situations in which blood-assisted devices are used are accompanied by high level of cardiovascular risk factors (high level of LDL, cholesterol, triglycerides and modified blood pressure).

214 Polyurethane

**Figure 8.** TAS evolution after PU/HPC incubation in blood plasma at 37 oC

Haemocompatibility involves compatibility with blood cells and blood plasma in other words, nonhaemolytic and nonthrombogenic behavior. Haemolysis is a mechanism by which erythrocytes (red blood cells) are destroyed through cell membrane lyses. Erythrocyte membrane lyses may occur as a result of environment pH modification or by cytotoxic action on erythrocyte membrane. Thus, both lipid (by lipid peroxidation) and protein (by

Thrombogenesis is a complex phenomenon by which thrombus is formed by blood clotting. As a physiological event, haemostasis implies the activation of the enzymatic cascades in which three main factors are involved – vascular, cellular and plasmatic (Edmunds, 1998).

A synthetic material can induce haemostasis activation by surface charge, hydrophobicity and/or released products of degradation. It is widely recognised that both positively and negatively charged as well as hydrophobic surfaces can induce thrombus formation. This

A positive charge can be favourable for FB adsorption, followed by its conformational modification and adhesion of platelets and leukocytes (monocytes). Adherent cells are activated and they release numerous molecules that lead finally to FB cleavage with fibrin network formation (clot). Among platelet secreted factors are platelet thromboplastin, fibrin stabilizing factor, serotonin, anti-heparin factor, and others. Adherent (activated) monocyte releases thrombogen tissue factor (TF). Mechanism triggered by positive and hydrophobic surfaces is mainly related to extrinsic coagulation pathway (B. Furie & B. C. Furie, 2008).

can be explained by involvement of several mechanisms as presented in Fig. 9.

*3.3.2. In vitro haemocompatibility* 

protein modification) compounds can be affected.

**Figure 9.** Contact coagulation and extrinsic (tissue factor) coagulation blood pathways

PUs are promising materials for implantable and non-implantable blood-interacting devices. They combine an increased elasticity with good mechanical resistance. For haemocompatibility evaluation of PU/HPC materials discussed above, the haemolytic and thrombotic potentials were determined by standard and adapted methods (see section 2). The obtained results are summarised in Table 2.


All studied membranes showed a low haemolytic activity, lower for PU/HPC than for pure PU-PEGAsample.

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of the Improved Polyurethane Membranes for Medical Applications 217

created, the homeostatic mechanisms are disturbed, and the resolution of the inflammatory response (simple put, the measure of the host reaction). All of these will finally establish the effective compatibility of a specific device. In the same time, understanding these reactions (the implant versus the host and the host versus the implanted device) will reduce health problems to the beneficiary of the device and device malfunction. Usually, for practical reasons, the homeostatic mechanisms are separately assessed even if it is well known that

The first event after a device/material insertion is that the body generates quickly a sort of "interface" *via* nonspecific adsorption of plasma/tissue soluble proteins on the implant surface (Shen & Horbett, 2001). There are some well identified elements that determine the FBR strength: device material composition, surface chemistry, size and shape, porosity,

As presented shortly below, tissue injury associated with device implantation, initiates a complex set of events (nonspecific inflammatory reaction and wound healing responses) that will bring about a FBR (Wahl et. al., 1989). The stages of inflammatory responses are

The initial phase, acute stage, starts quickly in matter of hours, lasts for several days (up to 14 days) and is underlined by rapid device interface generation and typical for this phase, different degree of neutrophil leucocytes responses (Jiang et. al., 2007). The main result of this stage is the building of temporary interface material-tissue, the cleaning-up of the injury

The acute inflammatory reaction typically decline in maximum 14 days with a "biocompatible" material. Some local conditions (extent of surgical injury, body reactivity) or properties of the implanted device can trigger a chronic inflammatory evolution

Numerous blood and tissue proteins such as cytokines (e.g. tumor necrosis factor (TNF), interleukins (IL-6, IL-8), matrix metalloproteases (MMP-1, MMP-3), granulocytemacrophage growth factors (GM-CSF)) are released, and leukocytes adhere to the endothelium of the blood vessels and infiltrate the injury site. These proteins are strong calling factors for monocytes, cells which will migrate to the site of inflammation where they will differentiate into macrophage. If inflammatory stimuli persist, the conditions that can lead to chronic inflammation are created. Cell population of this stage of inflammatory reaction is usually characterized by the presence of monocytes, macrophages, and lymphocytes (Bhardwaj et al., 2010). Also, in this step it can be noticed that the proliferation of blood vessels (angiogenesis), and connective tissue occurs that participate in remodelling of the affected area. The formation of blood vessels is crucial for wound healing, supplying necessary factors for tissues reconstruction. In the end, the granulomatous tissue is replaced by an extracellular matrix (ECM) that acts not only as a physical scaffold but also as an essential modulator of the biological processes, including differentiation, development, regeneration, repair, as well as tumour progression. The end phase of the FBR draws in wrapping the implant by a collagenic fibrous capsule that limits the implant and therefore

degradation, velocity as well as the place of device insertion (Ratner & Bryant, 2004)

well studied and can be separate in acute and chronic inflammatory periods.

place and the vasodilation that bring more blood in the affected area.

(Kirkpatrick et al., 1998).

they are profoundly interrelated (Sieminski & Gooch, 2000).

**Table 2.** Haemolytic and thrombotic potential of the PU/HPC samples: Hb-haemoglobin; FBfibrinogen; PT- prothrombin time

As for thrombotic action, a correlation between adsorbed FB, platelet adhesion and amount of formed clot was registered, while no significant variation was recorded for PT. This latter parameter was kept within the normal limits (see footnote 3).

The judgement strictly based on the haemocompatibility results permits to state that all examined materials have an acceptable thrombotic potential (referring to physiological requirements). Considering clot amount and all the other characteristics discussed above, it is obvious that PU-PEGA and PU-PPG/HPC are not suitable for long-time functional integration.

#### *3.3.3. In vivo biocompatibility and performance*

The technological progress achieved in the last decades in apparently unrelated areas (biomaterials, biotechnology, cell and molecular biology, tissue engineering, and polymer science) has generated a boost in the development and use of devices for medical and/or other type of applications (e.g. artificial organs, biosensors, catheters, heart valves) (Shastri, 2003). In spite of real improvement of this sort of devices there are still some important problems to face since implanted medical devices usually reveal different degree of loss of functionality over time after insertion (Göpferich, 1996). Tissue or blood-device interface interactions or a lack of biocompatibility resulting from the normal homeostatic response of the body to the implantation injury, determining an inadequate *in vivo* functionality and longevity, remains a serious concern (Callahan & Natale, 2008; Fujimoto et. al., 2007; Morais et. al., 2010).

In order to protect the body from the foreign object, under normal physiological conditions, the body reacts by several nonspecific mechanisms (immune and inflammatory cells recruitment), usually termed foreign body reaction (FBR) (Anderson, 2001). There is an imperative call for knowing the degree to which the pathophysiological conditions are

<sup>1</sup> Percentage of released Hb over negative control

<sup>2</sup> FB concentration remained in blood plasma after incubation. FB control was 2,98 ± 0,04 mg/ml

<sup>3</sup> Physiological normal value according to related laboratory are between 8,3 s and 11,3 s

<sup>4</sup> Percentage of blood clotting over negative control (blood without incubated material)

created, the homeostatic mechanisms are disturbed, and the resolution of the inflammatory response (simple put, the measure of the host reaction). All of these will finally establish the effective compatibility of a specific device. In the same time, understanding these reactions (the implant versus the host and the host versus the implanted device) will reduce health problems to the beneficiary of the device and device malfunction. Usually, for practical reasons, the homeostatic mechanisms are separately assessed even if it is well known that they are profoundly interrelated (Sieminski & Gooch, 2000).

216 Polyurethane

PU-PEGAsample.

Haemolytic

Released Hb

parameter was kept within the normal limits (see footnote 3).

*3.3.3. In vivo biocompatibility and performance* 

1 Percentage of released Hb over negative control

Material samples

fibrinogen; PT- prothrombin time

All studied membranes showed a low haemolytic activity, lower for PU/HPC than for pure

potential Thrombotic potential

PU-PEGA 6,7±0,2 2,79±0,04 11,06±0,4 1,40±0,08 40% PU-PEGA/HPC 5,2±0,1 2,87±0,04 10,9±0,09 0,82±0,05 29% PU-PTHF/HPC 4,2±0,2 2,90±0,01 10,9±0,09 0,86±0,05 15% PU-PPG/HPC 5,5±0,1 2,77±0,07 10,9±0,07 1,25±0,09 89%

As for thrombotic action, a correlation between adsorbed FB, platelet adhesion and amount of formed clot was registered, while no significant variation was recorded for PT. This latter

The judgement strictly based on the haemocompatibility results permits to state that all examined materials have an acceptable thrombotic potential (referring to physiological requirements). Considering clot amount and all the other characteristics discussed above, it is obvious that PU-PEGA and PU-PPG/HPC are not suitable for long-time functional integration.

The technological progress achieved in the last decades in apparently unrelated areas (biomaterials, biotechnology, cell and molecular biology, tissue engineering, and polymer science) has generated a boost in the development and use of devices for medical and/or other type of applications (e.g. artificial organs, biosensors, catheters, heart valves) (Shastri, 2003). In spite of real improvement of this sort of devices there are still some important problems to face since implanted medical devices usually reveal different degree of loss of functionality over time after insertion (Göpferich, 1996). Tissue or blood-device interface interactions or a lack of biocompatibility resulting from the normal homeostatic response of the body to the implantation injury, determining an inadequate *in vivo* functionality and longevity, remains a

In order to protect the body from the foreign object, under normal physiological conditions, the body reacts by several nonspecific mechanisms (immune and inflammatory cells recruitment), usually termed foreign body reaction (FBR) (Anderson, 2001). There is an imperative call for knowing the degree to which the pathophysiological conditions are

serious concern (Callahan & Natale, 2008; Fujimoto et. al., 2007; Morais et. al., 2010).

2 FB concentration remained in blood plasma after incubation. FB control was 2,98 ± 0,04 mg/ml 3 Physiological normal value according to related laboratory are between 8,3 s and 11,3 s 4 Percentage of blood clotting over negative control (blood without incubated material)

**Table 2.** Haemolytic and thrombotic potential of the PU/HPC samples: Hb-haemoglobin; FB-

(%)1 FB (mg/ml)2 PT (s)3 Adhered platelet

(cells x 105/mm2)

% blood clot amount4

> The first event after a device/material insertion is that the body generates quickly a sort of "interface" *via* nonspecific adsorption of plasma/tissue soluble proteins on the implant surface (Shen & Horbett, 2001). There are some well identified elements that determine the FBR strength: device material composition, surface chemistry, size and shape, porosity, degradation, velocity as well as the place of device insertion (Ratner & Bryant, 2004)

> As presented shortly below, tissue injury associated with device implantation, initiates a complex set of events (nonspecific inflammatory reaction and wound healing responses) that will bring about a FBR (Wahl et. al., 1989). The stages of inflammatory responses are well studied and can be separate in acute and chronic inflammatory periods.

> The initial phase, acute stage, starts quickly in matter of hours, lasts for several days (up to 14 days) and is underlined by rapid device interface generation and typical for this phase, different degree of neutrophil leucocytes responses (Jiang et. al., 2007). The main result of this stage is the building of temporary interface material-tissue, the cleaning-up of the injury place and the vasodilation that bring more blood in the affected area.

> The acute inflammatory reaction typically decline in maximum 14 days with a "biocompatible" material. Some local conditions (extent of surgical injury, body reactivity) or properties of the implanted device can trigger a chronic inflammatory evolution (Kirkpatrick et al., 1998).

> Numerous blood and tissue proteins such as cytokines (e.g. tumor necrosis factor (TNF), interleukins (IL-6, IL-8), matrix metalloproteases (MMP-1, MMP-3), granulocytemacrophage growth factors (GM-CSF)) are released, and leukocytes adhere to the endothelium of the blood vessels and infiltrate the injury site. These proteins are strong calling factors for monocytes, cells which will migrate to the site of inflammation where they will differentiate into macrophage. If inflammatory stimuli persist, the conditions that can lead to chronic inflammation are created. Cell population of this stage of inflammatory reaction is usually characterized by the presence of monocytes, macrophages, and lymphocytes (Bhardwaj et al., 2010). Also, in this step it can be noticed that the proliferation of blood vessels (angiogenesis), and connective tissue occurs that participate in remodelling of the affected area. The formation of blood vessels is crucial for wound healing, supplying necessary factors for tissues reconstruction. In the end, the granulomatous tissue is replaced by an extracellular matrix (ECM) that acts not only as a physical scaffold but also as an essential modulator of the biological processes, including differentiation, development, regeneration, repair, as well as tumour progression. The end phase of the FBR draws in wrapping the implant by a collagenic fibrous capsule that limits the implant and therefore

prevents it from interacting with the surrounding tissue. The main tissue events of the material-tissue interaction and wound healing are schematically presented in Fig. 10.

Biocompatibility and Biological Performance

of the Improved Polyurethane Membranes for Medical Applications 219

Unwashed (unpurified) material was implanted first, to highlight the importance of the properly prepared biomaterial for medical usage. From Fig. 11, A-A2 images, it can be easily seen as an intense acute inflammation reaction with numerous neutrophil polymorphonucleate leucocytes (PMN), edema and early fibrin network formation away from implantation site. These results suggest that an inappropriately prepared material at some stage in manufacture and/or manipulation can delay wound healing. As we expected, at 30 days of implantation (Fig.11, B-B2 images), inflammatory chronic reaction was really strong for related material, with the characters of neovascularised granulomatous tissue

In the end of this chapter, comparative study concerning long-time potential functionality based on evolution of chronic inflammation of PUs/HPC discussed above was done. The histological images of 30 days implanted, properly purified PUs/HPC are shown in Fig. 12. There were found chronic inflammations with VNGT and FBR with GC for PU-PEGA/HPC (A-A2 images) and PU-PPG/HPC samples (C-C2 images). Moreover, granuloma formation (G) as result of macrophage material degradation was present at material-tissue interface of

**Figure 12.** Light microscopy images of tissue response following 30 days subcutaneous implantation of washed PUs/HPC. A–A2, PU-PEGA/HPC; B-B2, PU-PTHF/HPC and C-C2, PU-PPG/HPC. B2, Masson's trchrome staining; all other images - HE staining. Objective magnifications are indicated in left bottom

(VNGT) and giant cells (GC).

corner

PU-PPG/HPC (C and C1 images in Fig. 12).

Morphologic aspects (light microscopy) of the acute tissue reaction to subcutaneous implanted polyurethane (PU-PTHF/HPC) at 10 days of implantation and chronic inflammation at 30 days of implantation are shown in Fig. 11. The study was conducted on Wistar male rats using the protocol described in section 2.

**Figure 10.** Fibrosis and fibrous encapsulation. End stage healing response to biomaterials. GF – growth factor (PD – platelet derived, T – transforming, bF – basic fibroblastic); IL – interleukin; PGL – prostaglandin.

**Figure 11.** Light microscopy images of tissue response to implantation of subcutaneous non-washed PU-PTHF/HPC (PU): A - A2, 10 days of implantation; B -B2, 30 days of implantation. All images – HE staining. Objective magnifications are indicated in the left bottom corner

Unwashed (unpurified) material was implanted first, to highlight the importance of the properly prepared biomaterial for medical usage. From Fig. 11, A-A2 images, it can be easily seen as an intense acute inflammation reaction with numerous neutrophil polymorphonucleate leucocytes (PMN), edema and early fibrin network formation away from implantation site. These results suggest that an inappropriately prepared material at some stage in manufacture and/or manipulation can delay wound healing. As we expected, at 30 days of implantation (Fig.11, B-B2 images), inflammatory chronic reaction was really strong for related material, with the characters of neovascularised granulomatous tissue (VNGT) and giant cells (GC).

218 Polyurethane

prostaglandin.

prevents it from interacting with the surrounding tissue. The main tissue events of the

Morphologic aspects (light microscopy) of the acute tissue reaction to subcutaneous implanted polyurethane (PU-PTHF/HPC) at 10 days of implantation and chronic inflammation at 30 days of implantation are shown in Fig. 11. The study was conducted on

**Figure 10.** Fibrosis and fibrous encapsulation. End stage healing response to biomaterials. GF – growth

**Figure 11.** Light microscopy images of tissue response to implantation of subcutaneous non-washed PU-PTHF/HPC (PU): A - A2, 10 days of implantation; B -B2, 30 days of implantation. All images – HE

staining. Objective magnifications are indicated in the left bottom corner

factor (PD – platelet derived, T – transforming, bF – basic fibroblastic); IL – interleukin; PGL –

material-tissue interaction and wound healing are schematically presented in Fig. 10.

Wistar male rats using the protocol described in section 2.

In the end of this chapter, comparative study concerning long-time potential functionality based on evolution of chronic inflammation of PUs/HPC discussed above was done. The histological images of 30 days implanted, properly purified PUs/HPC are shown in Fig. 12.

There were found chronic inflammations with VNGT and FBR with GC for PU-PEGA/HPC (A-A2 images) and PU-PPG/HPC samples (C-C2 images). Moreover, granuloma formation (G) as result of macrophage material degradation was present at material-tissue interface of PU-PPG/HPC (C and C1 images in Fig. 12).

**Figure 12.** Light microscopy images of tissue response following 30 days subcutaneous implantation of washed PUs/HPC. A–A2, PU-PEGA/HPC; B-B2, PU-PTHF/HPC and C-C2, PU-PPG/HPC. B2, Masson's trchrome staining; all other images - HE staining. Objective magnifications are indicated in left bottom corner

The absence of GC and rich granulomatous tissue ingrowth through large material pores was observed for PU-PTHF/HPC sample (Fig.12, B-B2 images). Morphological aspect for PU-PTHF/HPC implant suggests a material-tissue integration and regenerative remodelling. Moreover, the fibroblast-rich tissue ingrowth only from one side of the membrane highlights the bifacial behavior of the implanted sample, with potentially tubular or cavity-like device performances. Thus, considering PU-PTHF/HPC increased haemocompatibility, oxidative and other biocompatibility advantages discussed above, we presume a cardiovasculardevice performance for this PU sample.

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of the Improved Polyurethane Membranes for Medical Applications 221

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No.2, (July 2003), pp.580–589, ISSN 0021-9797

No. 1, (April 2010), pp. 54–60, ISSN 0021-9797

304-315, ISSN 0959-8103

2101, ISSN 1759-9954

78–86, ISSN 0378-5173

ISSN 0927-7765

1999), pp. 233-234, ISSN 0892-6638

**5. References** 

8424
