development and characterization fibrin and hyaluronan ......development and characterization of...
TRANSCRIPT
Development and Characterization of Fibrin and
Hyaluronan Coated Biodegrada ble Polyurethane Films
Joanna Dawn Fromstein
A thesis submitted in conformity wi th the requirements for the degree of
Master of Applied Science and Engineering Department of Chemical Engineering and Applied Chemistry
University of Toronto
"opyright by Joanna Dawn Fromstein 2001
National Library of Canada
Bibliothèque nationale du Canada
Acquisitions and Acquisitions et Bibliographic Services services bibliographiques
395 Wetlinglon Street 395. rue Wetlingtari OttawaON K1AON4 Ottawa ON K I A ON4 Canada Canada
The author has granted a non- L'auteur a accordé une Licence non exclusive licence allowing the exclusive permettant à la National Libmy of Canada to Bibliothèque nationale du Canada de reproduce, loan, distribute or sell reproduire, prêter, distribuer ou copies of this thesis in microform, vendre des copies de cette thèse sous paper or electronic formats. la forme de microficheffilm, de
reproduction sur papier ou sur format électronique.
The author retains ownership of the L'auteur conserve la propriété du copyright in this thesis. Neither the droit d'auteur qui protège cette thèse. thesis nor substantial extracts fiom it Ni la thèse ni des extraits substantiels may be printed or otherwise de celle-ci ne doivent être imprimés reproduced without the author's ou autrement reproduits sans son permission. autorisation.
Development and Characterization of Fibrin and Hyaluronan Coated Biodegradable Polyurethane Films
Joanna Fromstein M.A.Sc. Thesis, 2001
Department of Chernical Engineering and Applied Chemistry University of Toronto
A new family of biodegradable elastomeric polyurethane blends was created.
These novel materials were coated with thin layers of fibrin (and hyaluronan
(HA)). The effects of varying blend composition on the morpholosy, mechanical
properties and degradation rate were assessed. The polyurethane blends, which
degrade on the order of months to years, were found to have mechanical
properties similar to those of a commercially available polyurethane wound
dressing. Both fibrin and fibrinIHA coatings were achieved on the blends. The
surface topology of the polyurethane surfaces appeared to affect coating
cohesiveness. This was particularly true for the fibrin coatings, which
demonstrated a less uniforrn appearance than the fibrin1HA coatinps. The bi-
laminar wound dressings developed in this work could be particularly applicable
for skin graft donor site applications, where speed and quality of healin~; are
extremely important factors.
Acknowledgements
The author would like to thank Professor Kim Woodhouse for al1 of her support and
suidance throushout the course of this research. In addition, I would like to thank
Bob Chern~kay, Zuzanna Eperjesi and Tracy Leung for providinçj me with an extra
pair of hands; Drs. Fish and Merguerian for qualitative assessment of the blends;
and Lucie Martineau and Luwin Chiaco from the Defense and Civil lnstitute of
Environmental Medicine (DCIEM), Downsview, for their expertise and aid with the
in vivo model. I would also Like to thank the DCIEM animal care facility and Pane,
Shek from DCIEM for allowing me to perform the in vivo work on their premises.
I am greatly in debt to my mentor Christine Bense who helped to get my started
when I first joined the Lab group, and for al1 her useful organizational advice she
gave me to help me achieve the goals I set out for myself. Gary Skarja played a
fundamental role in keeping me on track, and ensuring that my thought processes
were sound. His wisdom, aloofness, and guidance have helped me tremendously in
times of confusion and stress. I would also like to thank the other Woodhouse
group rnembers who have been there for advice or just a friendly conversation
when I've needed it: Sandra Elliott, Allison Brown, Alexis Amour, Rebecca Smith,
and Catherine Bellingham.
Finally, I'd like to thank al1 of my family and friends who have been there through
al1 my indecisiveness, stress, and excitement. The ones who've Listened to me go
on and on about my research even thouzh they often don't understand a word I'm
saying. More than anyone else, though, I'd like to thank Doug Sinclair for being an
unending source of Love, rnoose, support and inspiration.
Glossary of Terms
Angiogenesis: Formation of new blood vessels, also termed neovascularization.
Apoptosis: Programmed cell death. A regulated process, involving a series of well-defined morphological changes.
Chernotactic: Cells can be encouraged to rnove in response to chernical gradients of these agents.
Coagulation: The process of converting liquid blood into a solid clot.
Cytokine: A small, multifunctional secreted protein that can act as a signal between cells, often by binding to cell-surface receptors.
Cytotoxic: Harmful to cells. A substance or component that can cause damage to or destroy cells i s said to be cytotoxic.
Dalton (Da): A unit of measurement for molecular weight, based on the mass of a hydrogen atom. 1 Da = 1.66~1 g.
Extracellular Matrix (ECM): The relatively insoluble network of polysaccharides, fibrous proteins, and adhesive proteins that provide cells with structural support in tissues and can also affect the development and biochemical functions of the cells.
Extrinsic Pathway: The process of clotting factor interactions leading up to the formation of a biood clot, initiated by the normal wound healing response.
Exudate: A protein- and white cell-rich liquid secreted through the walls of in tact blood vessels during inflammation.
Fibroblast: A common connective tissue ce11 that secretes collagen and other components of the extracellular rnatrix.
Foreign Body Giant Cell: Large cells formed through the fusion of monocytes and macrophages in an attempt to phagocytose material.
Foreign Body Response: The tissue reaction to an implant, leading to the formation of a wall of cells and granulation tissue between the implant and the rest of the body.
Glycosaminoqlycan (GAG): A long, linear, polymer composed of repeating disaccharide (sugar) units. GAGS play an important structural role in the extracellular matrix.
Growth factors: A class of extraceMar polypeptides that are released by activated cells, and can initiate cell migration, differentiation and tissue remodelling.
Haernostasis: The termination of bleeding, caused by blood d o t formation and damaged blood vesse1 contraction.
lntrinsic Pathway: The process of clotting factor interactions that causes blood clot formation, initiated by contact with a foreign body.
Keratinocyte: The predominant ce11 type found in the epidermal Layer of the skin, named for their production of keratin filaments.
Langerhans cell: A type of ce11 found in the epidermal layer of the skin that recognizes, processes and presents antigens to other inflammatory cells.
Lymphocytes: Critical cells of the immune system, responsible for producing antibodies, attacking foreign cells and regulating the immune response. There are two types of lymphocytes: T cells and B cells.
Macrophage: A phagocytic ce11 that ingests and destroys microorganisms and other foreign material.
Mast cell: Large cells present in connective tissues that release chemicals during inflammation.
Monocyte: A type of white blood cell that serves to ingest foreign particles such as bacteria and tissue debris.
Myofibroblast: A special phenotype of fibroblasts, responsible for the contraction of a wound by means of actin filaments.
Occlusive: A descriptor used for wound dressings that prevent or minimize the amount of water vapour that can pass through the material.
Phagocytosis: The cellular process of destroying foreign particles through engulfment, digestion and chernical degradation.
Platelet: A blood ce11 important in the process of blood clotting. Also termed thrombocyte.
Polymorphonuclear leukocyte (PMN): A type of white blood cell, also called neutrophil, that plays a major role in the defence against bacterial and similar insults to the body.
Proteoglycan: A qoup of glycoproteins that are composed of a core protein to which one or more glycosaminoglycans i s attached.
T cell: A type of lymphocyte, derived from the thymus, which plays an important role in regulating the immune response.
Thrombus: A blood clot.
Table of Contents
............................................................................ Glossary of Terms iv ... List of Figures .............................................................................. v m
................................................................................. List of Tables x ........................................................................... List of Appendices xi
Chapter One: Introduction and Literature Review ......................m............ 1 .......................................................................... 1 . 0 Introduction 1
.............................................................................. 1.1 The Skin 2 1.2 Cu taneous Wound Healing ......................................................... 4
................................................. 1.2.1 Wound Repair in the Dermis 4 ........................................................... 1.2.2 Re-epithelialization 8
1 .2.3 Healing in the presence of Biomaterials ................................... 9 1 . 3 Wound Dressings .................................................................. 10
................................................. 1 . 3.1 The ldeal Wound Covering 10 ............................................................ 1 . 3.2 Current Dressings 11
..................................................................... 1.4 Polyurethanes 12 1.4.1 Polyurethane Chemistry .................................................... 12
................................................. 1 A.2 Elastomeric Polyurethanes 17 1.4.3 Polyurethane Degradation .................................................. 18
1.5 Fibrin ............................................................................... 20 1.5.1 Fibrin in Wound Repair ..................................................... 20 1.5.2 Exogenous Fibrin Use ........................................................ 21
1.6 Hyaluronan ......................................................................... 22 1.6.1 Role of Hyaluronan in Healing ............................................. 24 1.6.2 Fibrin . Hyahronan Interactions ........................................... 25
1.7 Potential Applications ............................................................ 26 1.7.1 Skin Graft Donor Sites ....................................................... 26
Chapter Two: Research Objectives ..................................................... 28 2.0 Objectives .......................................................................... 28
2.0.1 Short-Terrn Objectives ...................................................... 28 2.1 Motivations Behind the Research ............................................... 28
............ Chapter Three: Fabrication of Biodegradable Polyurethane Blends -30 3.0 Introduction ........................................................................ 30 3.1 Experimental Methods ............................................................ 31
3.1 . 1 Materials ...................................................................... 31 3.1.2 Fabrication of the Blends ................................................... 32 3.1.3 Btend Characterization ..................................................... 40
3.2 Results and Discussion ............................................................ 42 3.2.1 Nuclear Magnetic Resonance Spectroscopy .............................. 42 3.2.2 Differential Scanning Calorimetry ......................................... 48
3 . 2 . 3 Polarized Light Microscopy ................................................. 54 3.2.4 Mechanical Tensile Testing ................................................. 57
......................................................... 3.2.5 Degradation Results 61 .................................. Chapter Four: Development of the Fibrin Coating 68
...................................................................... 4.0 Introduction.. 68 ............................................................ 4.1 Experimental Methods 69
4.1.1 Materials ...................................................................... 69 4.1.2 Fibrin Coating Development ............................................... 69
........................................... 4.1.3 Characterization of the Coating 71 ............................................................ 4.2 Results and Discussion 71
....................................................... 4.2.1 Coating Optimization 71 4.2.2 Fluorescent Light Microscopy .............................................. 75 4.2.3 Scanning Electron Microscopy ............................................. 76
Chapter Five: Incorporation of Hyaluronan into the Coatings .................... 80 5.0 Introduction ........................................................................ 80 5.1 Experimen ta1 Methods ............................................................ 80
5.1.1 Materials ...................................................................... 80 5.1.2 Fibrin/Hyaluronan Coatins Development ................................ 81 5.1 . 3 Characterization of the Coatings .......................................... 81
5.2 Results and Discussion ............................................................ 81 5.2.1 Scanninsj Electron Microscopy ............................................. 81 5.2.2 Lia ht Microscopy ............................................................. 84
Chapter Six: Conciusions, Contributions and Recommendations ................. 87 6.0 Summary ........................................................................... 87 6.1 Contributions ...................................................................... 87 6.2 Conclusions ........................................................................ 87 6.3 Recommendations for Future Investigation ................................... 88
References ................................................................................... 90
List of Figures
Figure 1 : Figure 2: Figure 3: Figure 4: Fiqure 5: Figure 6 : Figure 7: Fiqure 8: Figure 9:
............................................................ The structure of skin 3 .......................................... The phases of dermal wound repair - 5
Segmented polyurethane structure ......................................... 13 The two-step polyurethane urea synthesis chemistry .................... 14 Dependence of specific volume on temperature .......................... 16 A typical stress vs . strain curve for an elastomeric material ............ 18
...................................................... The coagulation cascade 20 .................................................. The structure of hyaluronan 22
The chernical reaction involved in producing the phenylalanine based ................................................................ diester chain extender 33
Fiqure 10: The apparatus used for synthesising the L-phenylalanine based chain ................................................................................ extender 34
Figure 1 1 : Generic chemical structure of the polyurethanes utilised in this ................................................................................. research 35
...................................... Figure 12: The polyurethane synthesis reaction 36 Figure 13: The apparatus used for the polyurethane synthesis ..................... 39 Figure 14: Proton NMR Spectra of the parent polyurethanes ....................... 43
............. Figure 15: Representative NMR spectra of A) Blend 1 and B) Blend 3 47 Figure 16: Differential scanning calorimetry scan of Blend 1 ....................... 49 Figure 17: DSC scan of PhelPCL2000 ................................................... 50 Figure 18: DSC of PhelPE01000 ......................................................... 52 Figure 19: Polarized Light microscopy images of blends 2 and 4 at 80x
magnification ........................................................................... 55 Fiqure 20: Polarized Light micrographs of: a) Phe/PCL2000, b) blend 1 , c) blend
2 , d) blend 3 and e) blend 4 ........................................................ 56 Figure 21 : Ultimate tensile stress values for the polyurethane blends. pure
PhelPCL2000 and two commercially available polyurethane dressings ....... 59 Figure 22: Ultimate tensile strain values for the polyurethane blends,
Phe/ PCL2000, opsite' and ~egaderm' ............................................. 59 Figure 23: Young's Modulus values for the polyurethane blends, PhelPCL2000,
0psite"nd ~egaderm" ............................................................... 60 Figure 24: Typical tensile stress-strain curves for the four polyurethane blends . 60 Figure 25: Mass loss as a function of time for the degradable polyurethanes .... 62 Figure 26: SEMs of PhelPCL2000 following buffer degradation. at 1000x
magnification ........................................................................... 65 Figure 27: Fiqure 28: Figure 29: Fiqure 30: Fiqure 3 1 : Fiqure 32:
SEMs of blend 1 following degradation. at 1 OOOx masnification ...... 66 SEM images of degradation of blend 2 .................................... 66 Degradation of blend 3. as seen through SEM ............................ 67 Degradation SEM images of blend 4 ....................................... 67 The fibrin coating. developed by Rubens et a l ........................... 68 The new fibrin coating protocol ............................................ 70
viii
Figure 33: Fluorescent Light micrographs of Oregon Green labelled fibrin coatings. A) Blend 1, as seen at 5x magnification. B) PhelPCL2000, at 10x magnification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . *. . . 75
Figure 34: SEM of fibrin-coated blend 1, at 25 times magnification ....... ... . .... 76 Figure 35: SEM at 1000x magnification, showing a ridge in the fibrin coating on
blend 3 .................................................................................. 77 Figure 36: SEM'S of fibrin coated polyurethanes, 5000x magnification: A)
~egaderm', B) Phe/PCL2000, C) Blend 1, D) Blend 2 , E) Blend 3, and F) Blend 4 .. ... .. ... . . ...... . .. . . . . . . ........ . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 78
Figure 37: Fibrin coated blend 2 , seen at 500x magnification ...................... 79 Figure 38: FibrinlHA coatings on Tegaderm performed usine, the methods of A)
adding the fibrinogen and HA at the same time, and 6) adding the fibrinogen first ................................................................................... 82
Figure 39: SEMs of fibrin and hyaluronan- coated biodegradable polyurethanes: A ) Phe/PCL2000, B) blend 1, C) blend 2 , D) blend 3, and E) blend 4 ............. 83
Fiqure 40: SEM images of blend 3 coated with A: fibrin and hyaluronan, and 6: fibrin only ............................................................................... 84
Figure 41: Blend 1, coated with HA and Fluorescently labelled fibrin, at 5x magnification ..... ..... .... ..................................................... ... . .... 85
Table 1 : The characteristics of an ideal dressing ....................................... 11 ...... Table 2: Sources of the materials used in the biodegradable polyurethanes 32
Table 3: Compositions of the four polyurethane blends ............................. 38 Table 4: Thermal transition data obtained by DSC .................................... 48 Table 5: Sources of materials used in creation of the fibrin coating ............... 69 Table 6: Count per minute values obtained for the fibrin coating using the oriçjinal
.................................................................................. protocol 73 Table 7: Effect of different rinse protocols on PhelPCLZOOO radiolabelled
fibrinogen adsorption .................................................................. 73 Table 8: Effect of different rinse protocols on Blend 1 followine, '251-fibrinogen
............................................................................... adsorption 74 Table 9: Final counts per minute of various degradable polyurethanes coated with
............................................................... radiola belled fibrinogen 74
List of Appendices
Appendix One: NMR Spectra A1-1 .................................................................................. A ~ ~ e n d i x Two: DSC Spectra .................................................................................. A2-1
........................ Appendix Three: Development of three-dimensional foams A3-1 Appendix Four: Preliminary in vivo evaluation of the fibrin coating ....... A4- 1
Chapter One:
Introduction and Literature Review
1.0 Introduction Every year, millions of skin-related wounds are treated in emerpency rooms and
hospitals worldwide. On atl but the most superficial of wounds, a wound dressing
i s required to function as a temporary protective barrier. The type of material
selected wi l l depend on the severity and characteristics of the wound site. For
numerous cases, inctuding ulcers, burns and skin graft donor sites, occlusive fi lm
dressings are comrnonly used[l].
In addition to providing protection from the external environment, a dressing
should allow for rapid, effective healing, to take place. Exogenous application of
fibrin, an important protein to the repair process, has been shown to improve
healing rates under numerous conditions. In particular, fibrin appears to
encourage cellular infiltration and formation of new blood vessels. Another
extracellular component shown to affect the healing process i s the
glycosaminoglycan hyaluronan (HA). HA i s known to affect cell movernent, blood
vesse1 infiltration and scar formation.
In this investigation, a novel biodeyadable polyurethane f i lm was coated with
either fibrin or fibrin and hyaluronan. These fibrin(1HA) coated polyurethanes
were evaluated as potential wound dressings, possibly for skin graft donor site
applications. The coated materials were characterized and assessed extensively in
vitro, as were the polyurethanes used in the study.
1.1 The Skin The skin i s the largest organ of the body, spanning over 2 m2 on average in
adults[2, 31. Although the skin i s quite thin, having an average thickness of 1 to 2
mm. i t s primary role i s that of a barrier, protecting the body's highly organized
structure from physical, chemical or biological insults from the environment[3]. In
addition to this protective role, the skin acts as a sensory receptor and
temperature regulator[2]. If the integrity of large portions of the skin is lost due
to injury or illness, the defect must be rapidly covered and protected in order to
prevent major Loss of functionality, and death [4].
There are two principal layers in the skin: the epidermis and the dermis (Figure
1 ). The epidermis, the outermost layer, i s primarily cellular, and i s composed
predominantly of stratified keratinocytes[2, 31. There are four sub-layers of the
epidermis, distinguishable by the extent of cellular differentiation. The top layer
of cells, the stratum corneum (horny layer), is formed entirely of terminally
differentiated keratinocytes[2]. The following three layers, the strata
granulosum, spinosum and basale, contain decreasingly differentiated cells[2].
The dead cells of the stratum corneum are continually shed and replaced by new
cells supplied by the stratum granulosum[3]. For a new keratinocyte starting in
the basal layer, it takes approximately 4 weeks to miprate up to the horny layer,
undergo terminal differentiation and be removed[5].
The epidermis is predominantly responsible for the protective features of the skin.
The layers are organized such that the keratinocytes line up to forrn closely
packed columns, connected by desmosomes, which utilize keratin filaments, to
join adjacent cells[2]. The protein keratin helps protect the skin from foreign
substances[b]. In addition, the stratum corneum i s formed of several layers of
keratinocytes surrounded by intercellular lipids[Z, 71. This lipid-rich structure
prevents substances that are hydrophilic from entering the skin even though it is
well hydrated. While the physical structure o f skin is organized such that
microbial infiltration i s minimized, there are also immunological agents present in
the epidermis, including Langerhans cells, T cells, and cytokines[8].
Figure 1: The structure of skin. The two primary regions are the epidermis and the dermis. Each of these domains is separated into sub-layers. The epidermis has five distinct layers, while the dermis has two. The primary function of the epiderrnis is to provide protection, and the dermal portion is responsible for the tissue's structural support.
The dermis provides the undertying structure and support to the skin[9]. I t i s
composed primarily of connective tissue, and has two main layers: the papillary
dermis and the reticular dermis[3]. The papillary dermis, the region closest to the
epidermis, i s significantly thinner, and i s abundant in type III collagen[3]. The
reticular dermis, on the other hand, i s predominantly fibrillar type I collagen, and
represents the bulk of the dermis[3].
The bulk, density, elasticity and tensile strength of the dermis are supplied by
fibrous proteins, particularly collagen and elastin[2, 31. Proteoglycans and other
smaller matrix molecules are responsible for permitting ce11 migration and
regulatinsj the osmotic properties of skin[2, 31. Fibroblasts, macrophages and mast
cells are the predominant cellular residents of the dermis[2]. Certain
inflammatory cells can also be found in close proximity to dermal blood vessels[2].
In addition to i t s structural properties, the dermis houses sensory receptors and
temperature regulators, and supplies nutrients and oxygen to al1 cutaneous
cells[6].
The dermis and epidermis are connected to one another through the basement
membrane zone (BMZ)[3]. This region, which i s predominantly a flat sheet of type
IV collagen, uses keratin filaments and type VI1 collagen fibrils as anchors, and
allows for tight integration of the two skin layers[Z, 51.
1.2 Cutaneous WoundHeaiing When the skin's integrity is disrupted, a complex dermal healing process involving
numerous cell types, cytokines, and the extracellular matrix is initiated[4]. This
procedure i s commonly divided into three overlapping phases: inflammation,
proliferation and remodelling (Figure 2)[4, 1 O]. Epidermal wound healing, while
inter-related to events occurring within the dermis, follows an alternate time
course. Section 1.2.1, below, deals solely with the dermal repair process. A brief
summary of epidermal healing follows, including a brief description of i t s
relationship with dermal healing.
1.2.1 Wound Re~ai r in the Dermis
1.2.1.1 Inflammation The inflammatory phase begins immediately following the disruption of blood
vessels in the tissue, and usually lasts for up to 4 days[4, 11]. This stase i s
responsible for achieving haemostasis and cleaning out the wound bed in
preparation for the formation of new tissue. The first event to occur i s the
appearance of platelets in the wound si te[ l l ] . These cells become activated by
the exposed collagen as they enter the wound site, causing them to release
extracellular matrix proteins, particularly fibrinogen and fibronectin[l 0, 1 11. In
addition to these glycoproteins, several cytokines and growth factors are also
released[9]. Concurrent with platelet activation, coagulation i s ini tiated,
predomînantly through the extrinsic pathway (Figure 7, page 20)[1 O]. The
resulting clot is composed primarily of cross-linked fibrin and platelets[lO]. This
fibrin matrix acts as a 3-dimensional scaffold, enabling cells to migrate into the
wound site[9].
Mvoflbroblasts 1
I Coltagen I I REMODELING
1 Hyaluronan 1 1 Flbroblasts 1
1 P D G F 1
I b . F G F PROLlFERATlON
Macrophages 1 Monocvtes 1
Leukocytes (PMNs)
1 Fibrln I
-- P tatetets
_I I
T i m e ( d a y s p o s t w o u n d i n g )
M a j o r e v e n t s w i t h i n e a c h p h a s e K e y c c l l i ~ l a r p l a y c r s
E x t r ; i c v t l u I a r coni p o ~ i e r i t s prcscrit C y t o k ~ r i c s ir ivolved
1 y e a r
Figure 2: Tlie phases of dermal wound repair. This figure, adapted from Mast[lZ], indicates the key events of the healing process and highlights the itnportant celltilar and extracellular players.
The chernotactic factors released attract
polymorphonuclear Leukocytes) soon after
Leukocytes remain in the wound site for the
neutrophils (also known as PMNs or
thrombus formation[lO, 131. These
first few days, and are responsible for
killing and phagocytosine, bacteria and matrix proteins[lO, 131. Following these
events, macrophages and their precursors, monocytes, are attracted to the wound
bed by cytokines, such as transforming growth factor P (TGF-P), released by the
platelets and leukocytes[4, 1 O]. Macrophages are responsible for digesting,
phapocytosing and killinp pathogens, scavenging tissue debris, and destroying the
remaining neutrophils[lO]. Once this wound debridement has taken place, the
wound bed i s ready for the second, proliferative phase, of wound healing[lO, 131.
As macrophages begin to enter the wound site, they secrete basic fibroblast
growth factor (bFGF) in order to encourage the formation of new blood vessels
(neovascularization)[4]. Angiogenesis i s an important part of the wound healing
process since cells migratin9 into a wound receive the necessary nutrients from
newly formed blood vessels that foliow the cells into the site[l4].
1.2.1.2 Proliferation
The proliferative stage commonly lasts from approximately the 3rd to the 6th day of
healing[ll]. This phase is characterized by the appearance of fibroblasts in the
wound site and the deposition of granulation tissue[4]. The macrophages present
in the wound bed provide fibroblasts with a continuous source of necessary growth
factors[4]. In particular, TGF-Pl and platelet derived growth factor (PDGF) are
responsibte for stimulating fibroblasts at the wound edge, and encouragine, them
to proliferate into the wound site[4]. As the fibroblasts migrate into the wound.
they Lay down the extracellular matrix necessary for ce11 movement[4]. This new
provisional matrix is constructed primarily of fibrin, fibronectin and hyaluronic
acid (HA) [4].
Once fibroblasts have infiltrated the wound, they begin to replace the provisional
matrix with a collagen-rich matrix[4]. TGF-Pl i s one of the major factors that
stimulate this collagen deposition[4]. Blood supply i s essential during this stage of
repair. Thus, angiogenesis continues to be stimulated during proliferation by bFGF
and by vascular endothelial growth factor (VEGF), which is produced by migrating
epithelial cells[4].
1.2.1.3 Remodelling Remodelling of the granulation tissue begins roughly 6 days post wounding[l 11.
This next phase is the Longest, lasting weeks to months. I t should be noted,
however, that these times are approximate, and are based primarily on
macroscopic observation of the wound site. Often, remodelling of the unordered
collagen within the scar will continue after healing i s considered to be
"complete"[l O]. Tissue remodelling i s characterized by two major events: wound
contraction, and the replacement of disorganized collagen fibrils with more
structured, mature scar tissue[4].
Wound contraction i s undertaken by fibroblasts that have assumed a myofibroblast
phenotype[4]. These myofibroblasts, which usually appear during the second week
of healinp, use bundles of actin filaments, linked through both cell-cell and cell-
matrix attachrnents to achieve the contraction[4]. Several cytokines, including
TGF-01 and 932, and PDGF are necessary for encouraging the transition to the
myofibroblast phenotype and subsequent contraction[4]. The myofibroblasts use
their actin attachments to the matrix to pull their surroundings towards them,
thereby compacting the granulation tissue[l O].
Throughout the remodelling stage, epithelial cells, fibroblasts, macrophages and
leukocytes al1 secrete tissue collagenases[lO]. As the disorganized granulation
tissue is broken down, fibroblasts Lay down new collagen fibrils, and begin to
reorsanize the tissue, forming a scar[lO]. The total arnount of collagen in the
wound site continuously increases until approximately 2 to 3 weeks after
injury[lO]. Even after the maximum collagen density i s reached in the wound
bed, the collagen continues to be remodelled for up to a year[lO]. Once the
wound i s futly healed, however, the tissue's tensile strenpth i s never restored to
more than 80% of i t s original strength[l O].
Once healing has been completed, the hiph cell population is no Longer necessary,
and so most of the cells undergo apoptosis, resulting in a predominantly acellular
scar[4]. This programmed cell death also leads to the disintegration of many of
the newly formed blood vessels, since a higher supply of nutrients i s no Longer
required .
1.2.2 Re-e~ithelialization The epidermis begins to repair itself within hours of injury[4]. The epithelial celis
transform phenotype, resulting in the retraction of intracellular tonofilaments.
and the dissolution of desrnosomes and hemidesmosomes[4]. These newly
transformed cells form actin filaments that they can use to aid in their
migration[4]. Collagenase (matrix metalloproteinase-1 MMP-1) produced by the
epidermal cells cleaves the newly exposed dermal collagen, creating fragments
upon which the cells can migrate towards the centre of the wound[4]. Exposure to
bare collasen i s also thought to promote proliferation, resulting in a new supply of
epidermal cells to f i l l in the defect[4]. As the cells that were at the wound edges
migrate inwards, their neighbours follow until the wound bed has been completely
covered[4]. Cytokines released by the epidermal cells during this re-
epithelialization process are thouglht not only to promote epithelial ce11 migration
and proliferation, but also to stimulate angiogenesis in the dermal wound[4].
In order for full re-epithelialization to take place, the epidermal cells must have
an underlying dermis to migrate along. Therefore, in deep wounds, scar tissue
must first be Laid down to replace dermal defects before the keratinocytes are
able to migrate, a process that can take days to weeks[l5]. If the wound i s
superficial, however, re-epithelialization can begin almost immediately post-
wounding[l6]. Until re-epithelialization can be completed, the wound i s highly
susceptible to infection and other environmental stimuli, as the skin's natural
barrier is not present.
1.2.3 Healina in the Dresence of Biomaterials The process of healing can be greatly affected by the presence of a fore
material in the wound site. While a wound dressing is a biomaterial, it belongs
9
ign
; to
a separate category of material, as it does not get implanted within a tissue.
implantation can often lead to chronic inflammation, frustrated phagocytosis, and
capsule forrnation[l7]. The result i s an acellular scar forming a barrier around the
implant in order to protect cells from the foreign body. This response i s termed a
foreign body response, and i s orchestrated primarily by foreign body giant cells,
macrophages and fibroblasts.
The presence of a wound dressing in contact with a wound site will, however, s t i l l
play a role in the healing process. The exposure of the wound dressing material to
the cells first entering the wound bed wil l initiate the intrinsic pathway of the
coagulation cascade (Figure 7, page 20)[17]. This pathway, like the extrinsic
pathway, leads to the production of a fibrin do t . Cellular reactions to a wound
dressing can lead to delayed healing due to a prolonged inflammatory response. In
addition, a more acute response could result. Phagocytic cells such as
macrophages and foreign body giant cells may undergo frustrated phagocytosis, a
process whereby the cells attach thernselves to the foreign surface. and secrete
oxidative, including enzymes, hypochlorous acid and nitric oxide derivatives, in an
attempt to break the material down[l8]. Since small fragments of the material
may be broken down and possibly internalized by phagocytic cells, and cells wil l
be present at the surface of the material, it i s important that the wound dressing
is not cytotoxic.
While the presence of a biomaterial can affect the inflamrnatory actions occurring
within the wound site, it does not prevent deposition of collagen within the wound
bed since the material i s at the surface. An increase in inflammatory response can
actually help the healing process, as the wound may be cleared of debris and
prepared for new collagen deposition more quickly. The presence of a wound
dressing can also help to provide a desirable healin? environment, as described in
the following sections.
1.3 Wound Dressihgs A wound dressing acts primarily as a protective barrier, replacing the function of
the disrupted epidermis[l9]. A wide variety of materials, ranging from simple
tauze to highly complex materials composed of both natural and synthetic
components, can serve as wound dressings. This diversity i s necessary since the
particular covering used in a given application wil l depend on the type of wound,
and the degree of protection required.
One common type of covering used for wounds that do not require excessive
removal of exudate i s the occlusive wound dressing. Also termed semi-occlusive or
semi-permeable, these dressings prevent wounds from drying out by limiting the
amount of water vapour that can pass through the material[l]. Occlusive
dressings are often composed of thin polyurethane films with acrylic
adhesives[20]. Semi-permeable films have been shown to increase epidermal
healing rates as compared to untreated wounds[20-221. In addition, patients
treated with occlusive dressings appear to experience less pain than with more
traditional dressings such as tulle[21]. The dressing developed in this research
thesis falls under the category of occlusive wound coverings.
1.3.1 The Ideal Wound Coverinq There are many characteristics that would be desirable in an optimized wound
dressing. The goal of such materials i s to provide an environment in which rapid,
effective healing can take place[l]. The perfect covering would therefore
maximize the healing rate, while s t i l l producing acceptable functionality and
cosmetic appearance in the resulting scar. While these are often the most
discussed features, there are many other design parameters that should be
optimized as well. The criteria listed in Table 1, below, indicate the wide variety
of concerns that can be included in the evaluation of a prospective dressing. In
addition to these features, certain application-specific characteristics may be
desirable. An example of this woutd be the use of a biodegradable material, as
was used in this project.
Adheres rapidly to the wound O
Prevents dehydration O
Does not require frequent changes
Enhances quality of healing
Conforms to the wound
Encourages haemostasis O
Prevents bacterial infiltration O
Allows for evaporation of exudates
Minimires pain and discomfort
Accelerates healing
Quickens re-epithelialization
Eases application and removal
1s sterilisable
Possesses elastici ty
Is easy and inexpensive to manufacture
Table 1: The characteristics that a wound dressing must possess in order to become the ideal dressing[4, 22- 251.
1.3.2 Current Dressincis There are numerous commercially available wound dressings, including a number
of semi-permeable materials. An important limitation associated with the
currently available occlusive wound dressings, however, is their inability to adhere
to the wound bed[l9]. Therefore they can only be used on relatively srna11 sized
wounds. In addition, some semi-permeable films allow too much fluid to build up
under the dressing[l]. Through the use of a thin fibrin (and hyaluronan) coating,
adherence of the dressing to the wound is likely to be greatly improved. Fibrin
sealants, which have essentially the same composition as the fibrin coating used in
this study, are often used to heLp dressings and skin grafts adhere to wound
beds[26-281.
None of the currently available dressings adequately address al1 of the desired
properties Listed in Table 1. The dressins developed in this study dealt with many
of these concerns. In particular, adherence to the wound bed, haemostasis,
healing rate and scar quality may be affected. In addition, through the use of a
biodegradable material, the covering may not need to be removed, thereby
preventing a re-opening of the wound and possible increases in pain, healing rate,
and scar formation. Since the polyurethane film acts as a semi-permeable
dressing, dessication may be prevented, and pain will hopefully be minimized.
The material also possesses elasticity, and can be sterilized using gamma
radiation. As with any biomaterial, cost i s an important factor. Through the use
of synthetic materials for the substrate, and only a relatively thin coating of
biological components, the cost of the dressing used in this thesis could be kept
relatively Low, as compared wi th some of the more expensive, naturally- based
materials currently on the market. Thus, while the developed dressing does not
incorporate every feature desired in an ideal wound dressing, i t does address many
of the more important issues.
1.4 Pdyurethanes Polyurethanes are used in a wide range of biomedical applications because they
have very good mechanical properties, and their biocompatibility has been well
established[29]. Elastomeric properties are particularly important for wound
dressing applications. Segmented polyurethanes provide a material that i s flexible
yet strong, a highly desirable combination for wound dressings, as welL as other
tissue engineering constructs[29, 301.
1.4.1 Polvurethane Chemistw Segmented polyurethane copolymers are essentially composed of domains of hard
segment embedded in a soft segment matrix (Figure 3)[29]. This phase segregated
morphology appears to be responsible for the polyurethane's superior physical and
mechanical properties[29]. By varying the compositions of both segments, the
physical properties of the polymeric material c m be altered.
Figure 3: Segmented polyurethane structure. The material is characterized by domains of hard segment embedded in a matrix of soft segment. Adapted from Lamba et al., with permission[31].
The soft sesment i s usually composed of a polyester or polyether diol of Low to
moderate molecular weight, while the hard segment contains a diisocyanate and a
low molecular weight chain extender[29, 301. The chain extender i s usually a
diamine or diol. Synthesis of polyurethanes i s often perforrned as a two-step
condensation process, whereby the diisocyanate i s first reacted with the soft
segment di01 to form a prepolymer. It should be noted that though this i s a
condensation reaction, no water molecule i s actually etiminated during the
reaction, as it i s immediately re-incorporated into the product. Next, the chain
extender i s added, in order to Link the short prepolymer segments together. The
result i s either a polyurethane urea or polyurethane, depending on whether the
c h a h extender is a diamine or diol, respectively. In this study, a diamine chain
extender i s being utilized. The two-step reaction scheme for synthesis of a
polyurethane urea i s shown below in Figure 4. The synthesis must be carried out
under a dry, nitrogen atmosphere in order to prevent side reactions of the
diisocyanate with water. The temperature must also be controlled in order to
prevent unwanted side reactions between the reactants from occurring[29].
Step 1 : Prepolymer formation:
Step 2: Chain extension:
prepoiyrner + H2N-RH-NH2
Figure 4: The two-step polyurethane urea synthesis chemistry. R, R ' and R" represent the citntral groups of the isocyanate, diol, and chain extender respectively.
In order to predict and explain the behaviour of polymeric materials in vitro and in
vivo, a firm understanding of the materials' properties i s required. Synthetic
polymers have several levels of structure. The primary structure simply represents
the (linear) chemical composition of the material. Secondary and tertiary
structures apply to the conformation of individual chains in solution[31, 321. The
morpholopy of a polymer represents chains with ordered conformations that
become oraanized in to Larper scale structures when they crysta llize[32].
Amorphous regions have no long-range order, and are characterized by randomly
entangled polymer chains[31]. More commonly, morphology has to do with the
extent of crystallization and the distribution of crystalline and amorphous regions
within the material. Segmented polyurethanes are generally semi-crystalline, as
they can never crystallize fully[31]. Some polymers do not crystallize at all.
Hydrogen bonding plays an important rote in the morphoiogy of polymeric
materials. Within a polyurethane urea, the urethane groups and urea groups will
interact with themselves and each other to form hydrogen bonds. As the
seamented polyurethane separates into microdomains of hard and soft segment,
urea-urea and urethane-urethane bondin9 will be stronser than urethane-urea
bonds, since the domains wil l align themselves in such a manner that within-
segment bonding will be more prominent[33].
Glass transition temperature (T,) can be used to determine a material's ability to
crystallize (Figure 5). If a polymer i s cooled from a gas, beyond the boiling,
freezing and glass transition temperatures, quickly enough that crystallization i s
not allowed to occur prior to solidification, an amorphous, glassy solid i s
formed[32]. Thus, crystallization i s governed by kinetic factors, and can be
altered by manipulating these factors. The glass transition of a polymer can be
detected since, although there i s no discontinuity in volume, specific heat changes
at this point. This specific heat i s the second derivative of free energy, and can
be detected using differential scanning calorimetry (DSC).
With segmented polyurethanes, the hard and soft segments have separate glass
transition temperatures (T,). The soft segment is usually a low to mid molecular
weight (400-5000) polyol which has a Low T,[29]. This low glass transition
temperature i s responsible for the formation of a rubbery, elastic, amorphous
matrix[31]. The hard segment contains two separate components: a diisocyanate
and a low molecular weight chain extender[29]. Typically, the hard seçjments
have a high glass transition temperature and are therefore either crystalline or
semi-crystalline[34]. lntermolecular associations are strongest in closely packed
crystalline regions, thereby making them harder, stronger and more chernically
resistant than their amorphous counterparts[31].
C; h s s
Figure 5: Dependence of specific volume on temperature. This figure indicates the ability to form a gtassy material rather than a crystalhne material i f the polymer i s cooled too quickIy, thereby preventing freezing from occurring prior to the glass transition temperature. Adapted from Painter and Coleman[32].
At low temperatures, the T, of the soft segment can affect the material's
mechanical and structural properties. As the temperature is elevated, the soft
segment becomes Less important, and the hard segment T, dominates. By
increasing the molecular weight of the soft segment, the polyurethane's glass
transition temperature i s decreased, leading to more extensive phase separation,
and less crystallization.
Polymers that have been heated to above their Tg gain enough enerçjy for rotation
about the backbone, enhancing mobility and flexibility[31]. This corresponds t o
melting of the crystalline material, and the temperature at which this occurs is
called the rnelting temperature (T,). This transition's Latent heat of rnelting can
be used to determine the percentage of crystalline material present in the
polymeric matierial.
Because material properties are so dependant on thermal properties, the thermal
history of a biomaterial is important and can affect the final result. Thus, it i s
important, when comparing materials, to ensure that they have been subjected to
identical thermal conditions.
1.4.2 Elastomeric Polvurethanes The morphological characteristics and thermal history of a material are important
in determining its mechanical properties. Tensile testing can yield valuable
information about the strength and elasticity of a material. An elastomeric
material has a characteristic stress-strain curve (Figure 6). This curve
demonstrates a linear resion early on, representing reversible elasticity. In this
region, for a given applied stress, the material can be stretched by a proportional
amount. Once the force i s released, the material wil l return to i t s original form.
The slope of this region i s known as the Young's Modulus, a property used to
compare different materials' abilities to be stretched reversibly . Beyond the
linear region i s a local maximum, representing the yield stress and strain of the
material. Once the material i s stretched past this point, permanent deformation
wil l occur. Eventually, if enough stress i s applied, the material is stretched as
much as possible and the material breaks. This point i s known as the ultimate (or
maximum) tensile stresslstrain point. Through com parison with the mechanica:
properties of biological tissues, materials that will be able to withstand the
biological environment can be selected.
-
Strain (% Elongation)
Figure 6: A typical stress vs. strain curve for an elastomeric polyurethane.
1.4.3 Polvurethane Dearadation Polyurethanes are commonly used in biomedical applications because they are
known to be quite stable[35]. In reality, however, al1 polyurethanes can be
degraded, to an extent, when exposed to in vivo conditions. This appears to be
achieved primarily through hydrolysis and oxidation[35]. Some chernical and
enzymatic degradation also takes place[35]. Often, non-degradabili ty i s a highly
desired trait in biomaterials. Sornetimes, however, intentionally degradable
constructs are needed. In these cases, polyurethanes can be produced using
components which are especially susceptible to hydrolytic, oxidative or enzymatic
degradation.
In order to produce a degradable biomaterial, components of both the hard and
soft segments can be chosen such that they are more susceptible to one or more of
the mechanisms of degradation. Commonly, soft segments have been utilized to
incorporate hydrolysable components into biomedical polyurethanes. Ester
Linkages are particularly susceptible to oxidative deçyadation [3 51. Known
biodegradable polyester materials such as polylactide and E-poiycaprolactone have
been utilized frequently to impart hydrolytic sensitivity to polyurethanes[34].
To incorporate susceptibility into the hard segment, researchers must focus
primarily on degradable chain extenders[34]. Incorporation of degradable chain
extenders is highly desirable since at relatively Low temperatures, the soft
segment mechanical properties dominate the polyurethane's properties. By using
a degradable chain extender, soft segments can be changed in order to alter
mechanical properties, without having to affect the degradation properties of the
material. Amino acids can be incorporated into diamine chain extenders. The
result is a chain extender that i s not only susceptible to hydrolysis at the amide
and ester linkages, but i s also sensitive to enzymatic cleavape[34]. Choice of
amino acid can be tailored to match the enzymes known to be present in the
desired site of use. For example, chymotrypsin-like enzymes present in the skin
contain a phenylalanine residue in the active site[34]. Therefore, the use of a
phenylalanine-based chain extender could increase the enzymatic susceptibility of
a wound dressing material. In addition, since amino acids are naturally found in
the body, degradation products are Less likely to be toxic if they contain arnino
acids.
There i s a low number of commercially available diisocyanates. Thus, very l i t t le
attention has been placed on this component when developing degradable
polyurethanes. However, many of the traditionally used diisocyanates, such as
toluene diisocyanate and 4,4'-diphenylmethane diisocyanate, may decompose into
carcinogenic diamines[34, 351. Thus, the development of diisocyanates based on
amino acids has led to research into their use in biodegradable polyurethanes[34].
These new diisocyanates are not particularly sensitive to degradative attack, but
can be used as a safety measure to minimize the likelihood of producing
potentially harrnfut degradation products when synthesising intentionally
degradable biomaterials.
Skarja and Woodhouse[U] have developed a family of biodegrada ble
polyurethanes tha t utilizes an L-lysine derived diisocyanate (LDI ) and L-
phenylalanine based chain extender (Phe) in the hard segment, and hydrolysable
polyethylene oxide (PEO) or polycaprolactone (PCL) for the soft segment. The
result was a family of degradable polyurethane elastomers possessing a wide range
of material properties. These materials were used as the substrates in this
research into coated biomaterials.
1.5.1 Fibrin in Wound Re~a i r Fibrin i s a complex glycoprotein that plays a central role in the early wound
healins process. Fibrin i s laid down by fibroblasts, forming a provisional matrix
that allows for ce11 migration and proliferation in the wound site. It i s formed in
the coagulation pathway (Figure 7) through a reaction between fibrinogen and
thrornbin. Factor X l l l then cross-links the fibrin mesh, resulting in an insoluble
clot.
Pro t hrombin Thrombin i ! I S1IIri
i
i, F~cror l S Ca:- (.il:-
Figure 7: The coagulation cascade. The extrinsic and intrinsic pathways are initiated by different stimuli, but lead to a comrnon reaction between fibrinogen and thrornbin to form a fibrin dot. Adapted from: Hanson and Harker[l7].
1 <.;? rissur F;ictor
Factor \ ' I I I t'latclrts
Factor S NP- Factor S
i Fïcnlr
C o w m P.\ r w Pl;iicIc.ts
Factor SlIl
1.5.2 Exoaenous Fibrin Use Medical professionals have used fibrin in attempts to promote haernostasis and
wound healing since before World War 1[27, 361. A product termed "fibrin
sealant" has been used extensively in a vast range of applications, including:
fixation of skin grafts, minimization of bleeding during open heart surgery, fixation
of dental implants and closure of incisions[27, 371. The U S . Army and the Red
Cross recently developed a wound dressing based on dry fibrin sealant[26].
Use of fibrin in biomaterials and during surgery i s attractive since fibrin i s a
natural substance encountered during, the healing process. Fibrinolysis. the
degradation of the provisional matrix and i t s replacement with collagenous tissues.
i s a standard step in wound repair[l7, 381. Durins this stage, fibrin dots are
broken down enzyrnatically by plasmin[38]. Exogenous fibrin sealant will also be
degraded via this mechanism, making it a biodegradable material.
In addition, use of fibrin sealant in healing wounds has been shown to have several
beneficial effects. In particular, it has been shown to: increase the rates of both
granulation tissue formation and re-epithelialization, promote angiogenesis, aid in
haemostasis, and improve fixation of dressings or skin grafts to the site[27, 37,
391. Fibrin has also been shown to be an effective drus carrier, providing a slow
release mechanism corresponding to i t s f i brinolytic degradation[40].
The fibrin coating developed in this thesis acts similarly to a thin layer of fibrin
sealant. Thus, the beneficial effects known to be associated with fibrin sealants
will be incorporated into the dressing. This design i s meant to provide a
recognizable matrix to encourage rapid ce11 migration into the wound site, and to
stimulate the neovascuiarization necessary for ce11 proliferation and viability. In
addition, fibrin and hyaluronan are known to have specific interactions that were
also incorporated into the design of this dressing. These interactions are discussed
in detail in section 1.6.2, below.
1.6 Hyduronan Hyaluronan (HA, hyaluronic acid) is a glycosaminoglycan found in the extracellular
matrix (ECM). I t is a high molecular weight polysaccharide of alternating
glucuronic acid and N-acetyl glycosamine units (Fisure 8). There are often 10,000
or more repeat units present in a single hyaluronan molecule. The molecular
weight of HA can Vary from approximately lo4 through l o6 Daltons[41, 421. HA is
an extremely attractive natural polymer for use in biomaterials since its purified
form is non-immunogenic. The structure of HA does not Vary from one species to
the next[43]. Hyaluronan has properties that enable it to play both biological and
physical roles in vivo.
Glucuronic acid N-acetyl glycosamine
Figure 8: The structure of the repeating disaccharide unit present in hyaluronan. This unit has a rnolecular weight of approxirnately 400 Da. The purple hydrogen balls represent axial hydrogen atorns. Adapted frorn Hascail and Laurent (421.
HA can be produced by most cells present in the body[43], indicating the
important role it plays in the ECM. Upregulation of HA production i s evident
during the early phases of wound healing, primarily during inflammation and
proliferation[43]. These levels remain higher for longer in foetal wounds, Leading
many scientists to believe that HA promotes scarlesz healing similar to that seen in
the foetus[43]. Hyaluronan i s commonly found in most soft tissues, including the
skin[43]. In fact, the skin serves as a primary reservoir for HA, holding over 50% of
the body's supply[44].
Hyaluronan plays a significant structural rote in the €CM of soft tissues due to i t s
highly viscous nature[43, 451. This viscosity i s proportional to the molecular
weight of the hyaluronan present[46]. HA i s also very hygroscopic, and as such,
appears to play an important role in keeping the skin hydrated[44]. In aqueous
environments, HA expands, filling spaces with a large, viscous polymeric network.
This ability to f i l l spaces i s thought to help in tissue organization[46]. In addition,
hyaluronan's Large size and ability to f i l l spaces appears to retard the movement
of larger molecules such as proteins, tissue degrading enzymes, viruses and
bacteria into HA-rich regions[43, 461.
Under physiological conditions, HA i s fully ionized, thereby influencing ion fluxes
necessary for ce11 signaltinçj[43]. I t scavenges free radicals and acts as an
antioxidant, modifying the extracellular environment even further[43].
Hyaluronan i s a highly polar molecule, but it has the ability to interact with non-
polar structures, demonstrating a unique amphiphilic nature[46].
Throush al1 of these mechanisms, HA has the ability to significantly impact the
extracellular space on a purely physical level. Originally, scientists believed that
HA i s simply a non-interactive packins molecule used to f i l \ spaces in the ECM[46].
However, in addition to playing, an important physical role, hyaluronan i s now
known to moderate cell behaviour in several ways. These effects are highly
dependent on concentration and/or molecular weight of the HA present[47]. I t i s
not fully understood how or why the molecular weight affects the response, but
varying viscosities are one possible explanation[47]. As a result, one must be
careful when comparing studies performed by various groups as the molecular
weight used can result in profoundly different results.
HA binds specifically to several types of cells and to certain extracellular matrix
components. There are three main classes of ce11 surface receptors that bind HA:
CD44, RHAMM (receptor for hyaluronan mediated motility), and ICAM-1
(intracellular adhesion molecule-1)[43]. CD44 i s suspected to play a role in cell-
ce11 and cell-substrate adhesion, ce11 migration, proliferation and activation, and
HA uptake and degradation[43]. RHAMM appears to affect primarily ce11
locomotion, while ICAM-1 regulates the degradation of HA[43].
In addition to direct contact with cells, HA can also modify cellular activities
through other extracellular proteins. The primary hyaluronan binding matrix
proteins (HABPs) are aggrecan and link protein[48]. Several agqecan and link
molecules can bind to one HA molecule, forming a cluster of HA surrounded by
aggrecan and link protein molecules[48]. 60th aggrecan and link protein help to
bind HA to the proteinaceous extracellular matrix, anchoring it in place and
contributing to the overall structure of the ECM. When bound to either aggrecan
or link protein, HA becomes shorter, adapting a more coiled conformation[48].
This new conformation may alter the functions of HA. Hyaluronan can also bind to
other extracellular proteins, including fibronectin, hyaluronectin, versican,
collagen VI, and P-32[48]. These other proteins appear to affect HA distributions
and conformations within the extracellular space, thereby rnodulating its function
and activity.
1.6.1 Role of Hvaluronan in Healinq Several studies have been performed investigating the effects of HA on wound
healing[ll, 41-45, 49-57]. Trends that have been observed indicate that cell
migration and proliferation are enhanced in the presence of both high molecular
weight hyaluronan and i t s degradation products[43, 47, 53, 561. More rapid
deposition of granulation tissue has also been observed in the presence of
exogenous high molecular weight HA[%]. This higher rate of qanulation tissue
formation is likely a direct result of more rapid cellular infiltration. More collagen
i s also deposited in the presence of HA, resulting in a denser granulation
tissue[58].
The presence of hyaluronan in high concentrations significantly reduces the
contraction of collagen matrices by fibroblasts[59]. This rnay be due in part to the
higher quantities of collagen present. With higher collagen concentration, there
rnay be less room for contraction. Wound contraction has a direct impact on scar
formation. In foetal wounds, HA levels remain high for significantly longer, and
wound contraction never occurs[51]. Studies have also found that in keloid scars,
there is significantly Less HA present in the papillary dermis than in that of healthy
skin[51, 601. Thus, hyaluronan appears to have a direct impact on scar formation.
Hyaluronan also modulates angiogenesis. Its degradation products are responsible
for encouragine, neovascularization[61, 621. Thus, as the hyaluronan present in a
wound i s enzymatically degraded, a supply of nutrients to the newly formed tissue
i s encouraged.
Through the use of hyaluronan in this wound dressins's coating, it was expected
that increased cell migration and tissue deposition within the scaffold would be
encouraged. As the cells begin remodelling their environment, they wil l degrade
the HA, thereby promoting vesse1 formation and continuing to support cell
migration within the wound site.
1.6.2 Fibrin - Hvaluronan Interactions Fibrin and hyaluronan each have distinct properties that make them zood
candidates for use in a tissue regeneration scaffold. They also have an additional
synergistic effect on one another, thereby making the use of the two materials in
conjunction even more desirable. Several groups have investigated the effects of
HA on fibrin d o t formation[40, 53, 63-68]. Recently, Hayen et al. have
determined that the presence of hyaluronan during fibrin polymerization changes
the morphology of the resultant clot[53]. With HA present, a more porous, more
turbid d o t was formed. These resulting clots were shown to stimulate turnour cell
migration[53]. If the HA was added after fibrin d o t formation, then there
appeared to be no effect on ce11 motility[53]. This indicates that miprating cells
prefer a wider, thicker substrate rather than a mesh made out of thinner
polymeric strands.
This same research group has also published results indicating that in vitro fibrin
clot configuration can be correlated to the arnount of capillary ingrowth and
endothelial cell migration within the clots[66]. In these studies, effects of
macromolecules (including HA), pH, ionic strength and thrombin concentrations on
the clot configurations were investigated. The conclusions made by this group
indicated that the hyaluronan-rich clots positively influenced ce11 mipration, but
not capillary formation. This may be due to the relatively short time line of 24
hours investigated. 60th HA and fibrin promote angiogenesis through their
degradation products rather than their polyrnerized forms, and degradation of the
clots may not have begun during the experimental time frame.
Weigel et al. have determined that hyaluronan specifically binds fibrin(o5en) [65].
This may play a role in changing the confiçyration of clots that have been formed
with HA. Nehls et al.[66] formed clots using large macromolecules, comparable in
size to hyaluronan, that do not interact with fibrin and fibrinogen. The results
were less stable, thinner clots than the one formed in the presence of HA. The
stability of the fibrin/HA clot has also been demonstrated by Wadstrom et al., who
found that fibrin sealant could be used as a method to slow the delivery of sodium
hyaluronate to a wound site[40].
Both fibrin and hyaluronan on their own have been shown to encourage cellular
ingrowth into a wound site. However, there i s further evidence indicating that by
combining the two, additional benefits may be seen. In this work, fibrin was
polyrnerized onto the surface of the scaffolds in the presence of high MW
hyaluronan in an atternpt to provide additional stability and the potential for
further encouragement of cellular infiltration. This configuration is expected to
help improve and speed up the wound healing process. As the bioactive coating
undergoes degradation, neovascularization may also be encouraped.
1.7.1 Skin Graft Donor Sites Every year, more than 1.25 million people in the United States are treated with
burn injuries[4]. The coverage of excised burn wounds often causes the creation
of a second wound in the form of a skin graft donor site (SGDS). To date, the split
thickness skin graft is the most common method of closing up burn wounds'. SGDS
wounds can be quite large i f large portions of the body have been burned. Often.
there i s more pain and debilitation caused by SGDSs than by the original burns[25].
As a result, the treatments for SGDSs should aim to allow the wound to heal as
quickly and completely as possible while minimizing pain and discomfort[25]. In
many cases, donor sites need to be re-harvested numerous times in order to obtain
enough coverage. As a result, speed of repair can be critical. Clearly, there i s a
need for a wound dressing which might facilitate the rapid re-epithelialization of
SGDSs without compromising the quality of wound care.
To date, occlusive dressings such as opsite" (Smith + Nephew) and TegadermB
(3M) are most commonly used on SGDSs[l]. They allow for a rapid re-
epithelialization by keeping the wound bed moist, while protecting it from
bacterial infection[69]. The fibrin and hyaluronan coated polyurethane developed
in this research provides the same advantages of commercially available occlusive
dressings while also encouragins haemostasis and more rapid, effect
healing. Thus, it i s expected that this new biomaterial could potentia
skin graft donor site applications.
ive wound
Uy benef it
Chapter Two:
Research Objectives
2.0 Objectives
The long-term objective of this research i s to develop a bi-laminar wound dressins
composed of a biodegradable polyurethane substrate coated with a thin layer of
fibrin or fibrin and hyaluronan. The materials under development may be
particularly useful for application on skin graft donor sites (SGDSs).
2.0.1 Short-Term Obiectives
The work described in this thesis encompasses three short-term objectives:
1. Production and characterization of a family of biodegradable polyurethane blends.
2. Development of a stable fibrin coating on both biodegradable and commercially available polyurethanes.
3. Development and characterization of fibrin / hyaluronan coatings on degradable polyurethane blend substrates.
2.1 Motivations Behind the Research
The objectives pursued in this work were selected in order to determine the
feasibility of developing a new wound dressing that wil l improve the quality
and/or rate of healing in the wound site. Through the achievement of these
objectives, important knowledge could be obtained about: the formation of new
degradable materials with superior properties, the ability to coat biodegradable
materials that possess a range of chernical properties with fibrin and
fibrin/hyaluronan, and the effect of hyaluronan on the formation of a fibrin
coating. The ability to form and characterize a reproducible fibrin and hyaluronan
coated biodegradable wound dressing could lead to a highly desirable biomaterial.
The degradative properties of the dressing could prevent the need to remove
28
dressings, thereby re-opening the wound. The incorporation of fibrin could
increase the rate of healine, by providing a pre-fabricated scaffold for cellular
infiltration to the wound site. The addition of hyaluronan to the fibrin could play
numerous roles, including the stimulation of angiogenesis, im provins cellular
mobility, and reducing wound contraction and therefore scarring. The work
described in this thesis provides some of the materiat developrnent and
characterization necessary to achieve the long-term goal of this research.
Chapter Three:
Fabrication of Biodegrada ble Polyurethane Blends
3. O Introduction The family of biodegradable polyurethanes developed by Skarja and Woodhouse
possess a wide range of material properties[34]. In this thesis, new blends, based
on this farnily of polyurethanes, were created. Through the use of blends,
material properties such as strength and degradation rate can be optimized.
The polyurethanes used in this work contain either a polyethylene oxide (PEO) or
polycaprolactone (PCL) soft segment, and a phenylalanine based chah extender
(Phe) and lysine based diisocyanate (LDI) hard segment[34, 70, 711. These
components were chosen due to their favourable susceptibilities to enzyrnatic and
hydrolytic attack, and their capacity to degrade into non-toxic products. Both PCL
and PEO are known to undergo hydrolytic degradation. L-phenylalanine i s present
in the reactive site for chymotrypsin-like enzymes found in the skin. Therefore.
enzymatic degradation may be encouraged by i t s presence in the chain extender.
The diisocyanate was not chosen to enhance degradation, but was selected
because it degrades into what are suspected to be non-toxic degradation
products[34]. Prelirninary studies involving polymeric materials made using L-
lysine diisocyanate (LDI ) have indicated the production of only non -toxic
products[34].
The nomenclature used for these polyurethanes indicates the chain extender used,
followed by the soft segment type and molecular weight. For example, a
polyurethane, made with LDI and Phe, and containing, PEO with a rnolecular weight
of 600 in the soft segment would be represented by Phe/PE0600.
In this research, four blends were
and one of either Phe/PE01000 or
to impart strength to the blends.
created. Al1 of the blends contain PhelPCL2000
Phe/PE0600. The PCL2000 material was chosen
On their own, the PEO-based materials cannot
be cast inio thin films appropriate for a wound dressing. They are "gooey" and
adhesive in their pure state. However, they are highly hydrophilic and therefore
have higher water uptakes than the PCL materials. Therefore the addition of
PhelPEOs into the blends is expected to enhance hydrolytic degradation. Thus,
through the formation of blends, materials possessing the strength of PCL2000 and
the degradation abilities of the PEOs could be created. By varying the
compositions of the blends, mechanical properties and susceptibility to
deyadation could be manipulated.
For a wound dressing, it is important to have a material that wi l l not tear, and wil l
stretch over joints without deforming permanently. The degradation rate of a
degradable wound dressing must be known in order to suit the dressing to a given
application. In the case of skin graft donor sites, a degradation rate of
approximately two weeks would be ideal. In this portion of the thesis, the four
blends and pure PhelPCL2000 were characterized using mechanical tensile testing,
polarized light microscopy, NMR and buffer degradation studies in order to
determine which materials might be best suited for a wound dressing application.
3.1 Experimenta/ Methods
3.1.1 Materials
The materials used in the fabrication of the chain extender and polyurethanes are
listed in Table 2, below.
Item Source
Chloroform (ACS) t ACP Chernicals Inc., Montreai, i Québec l
1,4-Cyclohexane dimethanol Milwaukee, WI (CDM)
Dimethylformamide (DMF) Aldrich, Milwaukee, WI 1
2,6-Diisocyanato methyl Kyowa Hakko Kogyo Co. Ltd., caproate (LDI) 1 Tokyo, Japan
, L-Phenylalanine (Phe) 1 Sigma, St. Louis, MO 1
Item Source Polyethylene s$ycol, MW600
I Aldrich, Milwaukee, W1 ' (PEO) I
Polyet hylene glycol, MW 1 000 l
Aldrich, Milwaukee, WI
p-Toluene sulfonic acid Sigma, St. Louis, MO
(p-TsOH )
: Stannous-2-ethylhexanoate Sigma, St. Louis, MO
Table 2: Sources of the materials used in the fabrication and casting of the biodegradable polyurethanes.
3.1,2 Fabrication of the Blends Chain Extender Synfhesis
The L-phenylalanine based diester chah extender and al! three polyurethanes used
in these studies were synthesized following the methods of Skarja and Woodhouse
(Figure 9)[34]. First, the chain extender was synthesized, using a Fischer
esterification reaction between L-phenylalanine and 1,4-cyclohexane dimethanol.
The apparatus used in this synthesis i s illustrated in Figure 10, below. Initially,
700 mL of toluene, and approximately 70 8, of L-phenylalnine were combined in
the reaction flask. Next, approximately 80 g of p-toluene sulfonic acid was added
as a catalyst for the esterification reaction. The entire reaction vesse1 and the
water trap were then insulated using aluminium foi[, and the flask contents were
heated. 30 g of cyclohexane dimethanol (CDM) were then added to the reaction
flask, and the solution was allowed to heat until reflux began.
Figure 9: The chernical reaction involved in producing the phenylalanine based diester chain extender.
Once reflux had begun, the reaction was left to proceed for approximately two
hours. The presence of white solid at the bottom of the vessel indicated
completion. The apparatus was then allowed to cool, and the toluene was
decanted from the reaction vessel. The solid chain extender was broken into
small pieces, approximately 1 cm in diameter, collected in a beaker, and allowed
to dry at roorn temperature for at least 48 hours.
Next, the solid was dried in a vacuum oven at 80°C for 24 hours. Following drying,
the solid was ground up into a fine powder using a mortar and pestle. This process
was later optimized through the use of a porcelain ball rnill.
After drying, the powder was stirred in a large Erlenmeyer flask with
approximately 750 mL of ethanol for 2 to 3 hours, and then filtered usin-, vacuum
filtration. The powder retained by the f iker was re-dissolved in another 750 mL of
ethanol, and this procedure was repeated until the filtrate liquid emerged clear
rather than yellow. This step usuatly required 4 to 5 iterations. Once the filtrate
drying tub
, i
Cotton
0' dneri te
1 7
condenser -? l 1 water out
1 / water in 0 1
clamps ,_
round flask
~ h e a t i n q mantle
<
I - 1 -
i 1 u 1
1 ' l I 1 --
magnetic stir plate i l l heat controt
Figure 10: The apparatus used for synthesisins the L-phenylalanine based chain extender. This figure was adapted with permission from Sandra Elliott.
was clear, the solid was vacuum filtered usine, pyrogen
collected, and dried in a vacuum oven a t 80°C for 24 hours.
free distilled water,
The chain extender was then dissolved in pyrogen free water to make a 10% chain
extender solution (by weight). A molar excess (2x) of potassium carbonate was
then added to the solution, in order to neutralize any acidic residues remaining in
the chain extender. The chain extender and potassium carbonate solution was
heated and stirred until the solution began to clear. and a yellow solid remained.
The product was then vacuum filtered with distilled water, collected, and dried at
room temperature in a vacuum oven for 48 hours.
Finally, the chain extender was dissolved in approximately 250 mL chloroform, and
gravity filtered using Whatman grade 42 filter paper. The filtrate was then dried
at room temperature untii most of the chloroform had evaporated, approximateiy
two weeks. Next, the chain extender was dried in a vacuum oven at room
temperature overnight to ensure that al1 traces of chloroform had been
eliminated. NMR analysis was performed on the chain extender to ensure that the
chain extender matched previous batches in composition.
Polyurethane Synthesis:
The polyurethanes were produced usinp a two-step synthesis method, as outlined
in Figure 12, below. In addition to the three necessary components of the
polyurethane, stannous-2-ethylhexanoate was used as a catalyst during the pre-
polymerization. The ratio of reactants used was 1:1:2 for chain extender : soft
segment : diisocyanate. Thus, the expected repeat structure of the polymer was:
L
Chain Extender
Figure 1 f : Generic chernicai structure of the polyurethanes utilised in this research. The parts of the structure derived from each of the chain extender, diisocyanate and di01 are highlighted in biue, green and purple, respectivety.
LDI O PCL diol O PEO
P 7 O-C-N- CH-(CHz)4-NC0 +
c=o O
1 I
C h prepolymer I Phe chain extender
t
CH3 polyurethane urea
Figure 12: The polyurethane synthesis reaction. Since two different soft segments were used, they are represented from the prepolymer on as a squiggle. Figure adapted from Gary Skarja, with permission.
The polyurethane synthesis apparatus, shown in Figure 13, page 38, i s similar to
the set-up used for the chain extender. However, this reaction must be performed
under a dry nitrogen atmosphere in order to prevent side reactions with water
molecules. In order to produce the prepolymer, the apparatus was set up as
shown, and the nitrogen gas atmosphere was established. Next, approximately 8.5
g of LDI in approxirnately 180 mL DMF was added to the reaction vessel, and the
heating rnantle was turned on.
The necessary soft segment, calculated based on a 2:1 molar ratio with the LDI,
was heated and dried out in a vacuum oven at 60°C for 48 hours prior to synthesis.
Once the LDI was added, the dial was removed from the vacuum oven, re-
weighed, and dissolved in 90 mL of DMF. This mixture was then placed in the
addition funnel, where molecular sieves drew residual moisture out of the soft
segment.
Once the temperature inside the reaction vessel had reached 85"C, 0.1 mL of the
catalyst stannous 2-ethylhexanoate was added to the LDI solution. The soft
segment was then allowed to slowly drip into the reaction vessel over the course
of approximately 30 minutes. This pre-polymeriza tion reaction was then carried
out at approximately 90°C for 2.5 hours. Towards the end of the reaction, the
temperature control was turned off, allowing the vessel contents to begin coolinp.
As the prepolymerization reaction neared completion, approximately 8.8 g of L-
phenylalanine based chain extender were dissolved in 100 mL DMF. The chain
extender was then desiccated usiner fresh molecular sieves, and added to the
vessel once the reaction temperature had cooled to below 60°C.
The reaction was left to proceed overnight. In the morning, the contents of the
reaction vessel were poured into a saturated solution of potassium chloride in
pyrogen free distilled water, and allowed to s i t for 2-3 hours. The precipitate
formed was then removed from the solution and vacuum filtered using distilled
water.
The retentate was transferred to a beaker filled with distilled water, and
incubated at 37OC for 48 hours. The polyurethane was then vacuum filtered again,
and dried in a vacuum oven at 60°C for 2 days.
Table 3 indicates the compositions of the four blends investigated. In order to
simplify notation, the blends will be referred to as blends 1 throuph 4 elsewhere in
this thesis. To fabricate the blends, the parent components (PheIPCL2000,
Phe/PE0600 and Phe/PE01000) were first solvent cast in order to purify the
materials. Each polyurethane was dissolved in chloroform, at a concentration of
approximately 5% w/v, and gravity filtered usin5 Whatman 42 filter paper. Next,
the polyurethane-containing chloroform was poured into Teflon dishes, covered,
and allowed to evaporate for 48 hours. Once al1 the solvent had evaporated, the
polyurethane films were dried in a vacuum oven at room temperature for 48 hours
in order to remove any remaining solvent. These cast materials were then
weighed out in the corresponding ratios for each blend, and solvent cast followine,
the same method as described above. Al1 experiments performed in this research
utilised cast films prepared in this manner. In aU cases, ratios represent masses of
each polyurethane blended with the other prior to dissolution and filtration. The
final ratios of PEO-polyurethane to PhelPCUOOO are unknown, and tikely Vary
from one sample to the next. NMR analysis was attempted in order to investigate
this issue, as discussed in section 3.2.1.
P hePCL2000 PhelPEObOO Phe/PEO 1 000 Polyurethane Blend
(% by mass) (% by mass) (96 by mass)
Blend 1 75 25
Blend 2 75 25 .
Blend 3 50 50
Table 3: Compositions of the four polyurethane blends used for this thesis.
Originally, attempts were made to make blends containing more than 500
Phe/PE0600 or 1000. The results, however, were materials that did not maintain
their cast shapes and would therefore not be useful as a wound dressing. Upon
creation of blends 1 and 2, the materials were taken to two clinicians to see which
material they preferred based on feel. 00th doctors selected blend 2. the
material containing more PhelPE0600. The materials were also tested for
suturability. The results of this test indicated that blend 2 has a tendency to tear
when tension i s applied against sutures. The 75%: 25% blend was significantly
better at withholding force applied to sutures. Since the one material was
preferred by clinicians, and the other demonstrated good suturability , both
blends, and their PEOlOOO counterparts were carried throughout this research.
3.1.3 Blend Characterization
Nuciear Moqnetic Resonance Spectroscopy (NMR):
Two samples of each type of blend (from different batches), as well as of al1 three
parent polyurethanes, were sent for proton NMR analysis. The solvent used was
d6-dimethyl sulphoxide (DMSO). The samples were run on a 400MHz Bruker
spectrometer, with 64 scans. Peak areas and peak heights were included on the
spect ra.
Differentioi Scanning Colorimetry
As for the NMR analysis, two samples of each material were sent out for analysis
using differential scannine, calorimetry (DSC). Samples approximately 10 - 15 mg
in size were run, using a Thermal Analyst 2100 thermal analyser (TA Instruments.
Newcastle, DE). The temperature range used was -140°C to 250°C, with a ramping
rate of 15" per minute. The resulting plots of heat flow as a function of
temperature were obtained, and analysed for glass transition temperature and
crystallinity.
Polarized Light Microscopy :
Circular disks (8mm diameter), created from the cast films using a metal punch.
were examined under crossed polarized filters on a Lipht microscope (Axiovert
5100, Zeiss, Germany). Images were captured using Northern Eclipse v. 5.0
software (Empix lmaging Inc.)
Mechunical Tensile Testing:
Tensile testing was used to determine the mechanical properties of the blends.
Pure Phe/PCL2000 and two commercially available occlusive wound dressings,
~egaderm"3~) and opsite" (Smith + Nephew), were used as controls. A
traditional uniaxial dogbone tensile testins method was used, following the ASTM
standard D638M. The samples were conditioned at 2 3 O C and 50% relative humidity
for a minimum of 48 hours prior to testing. An lnstron testing machine (SINTECHI,
SINTECH, Stoughton, MA) was used, at a strain rate of 500mm/min, until failure.
TestWorks II, v. 2 .10~ software (SINTECH) was used to collect and senerate the
data. Force and displacement data was converted to stress and strain values by
the software. From this data, stress vs. strain curves were generated using
~ ic rosof t ' Excel (Microsoft Corp., Seattle, WA).
Degrada tion Studies:
In order to determine the degradation properties of the polyurethane blends and
the parent polymer Phe/PCL2000, circular disks, 8rnm in diameter, were made
using a metal punch and the cast films. The disks ranged in thickness from
approximately 40 to 60 pin. Four replicates of each material were made for each
time point. The disks were then washed in distilled water containing 1% Triton X-
100, and allowed to dry in a dessiccator for 48 hours. The degradation studies
were carried out in tris buffered saline (TBS, 0.05M Tris, 0.1M NaCl), a t a pH of
7.4. Each sample was weighed, placed in a separate 2OmL vial, and 13mL of TBS
was added. The vials were then incubated at 37OC, and samples were removed
following, 3 days, 1 week, 2 weeks, 4 weeks and 6 months. After removal, the
samples were rinsed three times in distilled deionized water and allowed to dry a t
room temperature for 1 week. The samples were then re-weighed in order to
determine mass loss, and set aside for scanning electron microscopy (SEM)
analysis. Four fresh samples of each polyurethane were also washed and set aside
at the start of the study as controls. Mass loss values were determined using the
following calcuLation. This formula provides total mass loss values, from time O to
each time point, rather than relative mass losses between time points:
(Initial 1lnss)- (Final M a s ) O MISS L.0ss = 1 l o o O o
(Initial Mass)
Once the study was completed, al l samples were prepared for SEM. First, al1 four
samples of each type were affixed ont0 metal mounts using double-sided electrical
tape. The edges of the samples were then coated with silver paint in order to
improve electrical conductivity. Finally, the mounted samples were sputtered
with a very thin layer of platinum. The samples were examined under a scanning
electron microscope (mode1 5-2500, Hitachi, Japan), and the images were
captured using a PCI (passive imape capture) system (Quartz PCI, Quartz,
Vancouver, B.C., Canada).
3.2 Resu/ts and Discussion
3.2.1 7 From the spectra of the pure Phe/PE0600, Phe/PCL2000 and Phe/PE01000 (Figure
14, below), the soft segment peaks were isolated frorn those that represent
elements of the hard segment, which i s the same in al1 of the materials. In total,
there are fourteen chemically different protons identified in Figure 14. The
spectra for PhelPE0600 and Phe/PE01000 both contained the same peaks, as they
have the same chemical structure. Thus, one was used to label soft segment. and
the other hard segment peaks.
Peak assignments were performed through analysis of the chemical structures to
determine how many neiphbouring protons wouid split each peak, and through the
use of tables found in organic chemistry and NMR reference books(72, 731. Rules
about the effects of neighbouring functional groups on the protons were also taken
into account in order to determine relative positions. Using this technique, peaks
and/or approximate positions, were assigned for al1 the protons. In some regions,
the peaks are rather close together, due to the large number of ethyl groups
present within the polyurethanes' structures. In these cases, more than one peak
may have been assigned to the same region of the spectra.
Figure 14: A) NMR spectrum of Phe/PE0600, with the PEO soft segment proton group labetled.
Figure 14: 0) NMR Spectrum of PheIPCL 2000 demonstrating the soft segment proton peaks.
Figure 14 (cont'd): Proton NMR Spectrurn of C) Phe/PEOf 000. The hard sesment proton peaks are tabelled on this spectrum since the soft segment peaks will be the same as those identified in 148. Protons on the benzene and cyclohexane rings will in fact have several peaks, dependin9 on orientation and position of the protons, but were treated as one group of chemicatly similar protons.
In al1 of the PEO polyurethanes, a pentet peak centred at a chemical shift of 3.51
pprn (+O031 represents the CH2 protons of the PEO. This peak is labelled as #l.
The PCLZOOO material contains three chemically unique protons within the soft
segment. These peaks were identified, and ratios between spectra compared to
ensure that one peak did not Vary more than another. A pentet stretching from
1.5 to 1.6 ppm, Labelled as # 3, represents the central CH2 groups in the PCL
structure (Figure 14). This peak was selected for comparisons between PEO and
PCL peaks present in the blends.
The reasons for choosing peak 3 for comparison with peak 1 in the blends was
based on the relatively unique placement of the Phe/PCL pentet at 1.5 pprn.
Many of the remaining 12 protons in the polyurethane have similar chemical
environments, and therefore similar chemical shifts on the NMR spectra. It was
important to choose peaks that were easy to differentiate. Analysis of polymeric
materials through proton NMR is always relatively difficult. When looking, at
blends where everything but the soft segment i s the same in the two materials
blended, overlappage and interference of chernically similar protons may
complicate the analysis even further.
Using these two selected peaks for comparisons, a definite trend was seen when
observing the blend spectra. The ratio of the PEO peak area to the PCL peak area
increased from the 75/25 blends to the 50150 blends. Al1 of the peak heights and
peak areas were calculated and compared in an attempt to quantitate the relative
amounts of PhelPCL and PhelPEO polyurethanes. However, no accurate
quantitative data differentiating the blends from one another could be
established. Thus, it remains unknown whether a 75/25 blend i s indeed 75%
PhelPCL2OOO and 25% Phe/PE0600 (or 1000). These spectra did confirm, however,
that the amount of PEO incorporated into the blend increases for blends 3 and 4.
Figure 15, below shows representative NMR spectra of a 75%: 25% blend, and a
50%: 50% blend. ALI of the spectra are available in Appendix One.
Figure 15: Representative NMR spectra of A) Blend 1 and 0) B1end 3, with the appropriate peak assignments. From these two spectra, the seneral trend that the PEO peaks get relativeiy Larger in the 50%: 50% blends is seen.
3.2.2 Differential Scannins Calorirnetrv
The results from the DSC spectra are summarized in Table 4, below. ALI of the DSC
spectra can be seen in Appendix Two. Previous work with the parent
polyurethanes by Skarja[34] showed that the hard segment does not crystallize,
likely due to the bulky side chains present in lysine diisocyanate. If there were
any hard segment crystallinity, however, it would be shown above 1 50°C, beyond
the region of the scans. Therefore, the endotherrns seen on the DSC spectra
represent primarily soft segment characteristics. As can be seen in Table 4, al1
four blends and Phe/PCL2000 demonstrate a phase mixed, semi-crystalline
morphology, as i s evidenced by the presence of approximately 15 to 35'
crystallinity .
Material TS ( O c ) T m ( O c ) 1 Heat of Crystallinity
Fusion (J/s) (%)
Phel PEObOO -4.8 NIA N/A N/A
Phe/PE01000 -33.5 N/A NIA NIA
Blend 1 -14.0 46.5 33.3 24.6
Blend 2 -32.9 45.6 30.1 22.2
Blend 3 -19.8 46.6 22.9 16.9
Blend 4 -37.7 47.0 25.4 18.7
Table 4: Thermal transition data obtained by DSC. The figures in bold itaIics represent values that are likely inaccurate. ALI values, except those for Phe/PE01000 are averages of spectra obtained frorn two sarnples.
There is a relatively large degree of error in this analysis, as i s represented by five
values in the table whose accuracies are uncertain. These inaccuracies are the
result of poor spectral resolution. For blend 1 and Phe/PCL2000 the glass
transition temperatures were difficult to pick out on at Least one of the spectra.
An exarnple of such a scan for blend 1 i s shown below in Figure 16. For the
Phe/PCL2000. the material was quenched and re-scanned. It was only this re-
scanned spectrum that indicated a glass transition at all. The value obtained
through this quenched sample does agree, however, with the previously published
results of Skarja and Woodhouse[34].
The other Phe/PCL2000 values Listed in Table 4 were tabulated using only the un-
quenched scans. The O crystallinity vaiue of Phe/PCL2000 i s also imprecise since
one of the scans demonstrated a dip prior to the meit endotherm (Fisure 17), so
the area under the curve i s probably not accurate. The heat of fusion for this scan
was based on a sum of the two separate values calculated for the dip and then the
spike.
:-B IST RUN
Figure 16: Differentiat scanning calorimetry scan of Blend 1, illustrating the difficulty in picking out &ss transition temperature for some scans.
Blend 2 demonstrated g l a s transitions on both spectra, but the values varied
significantly. A ma l t exothermic dip is seen after the Tg on one of the blend 2
spectra. similar to the one seen for Phe/PCL2000 as well. This dip may have
altered the actual Tg value somewhat. contributing to the discrepancy between
the two samples.
Fisure 17: DSC scan of Phe/PCU000, where a dip was seen irnmediateLy prior to the melt endotherm. causing problems in calculatin~ meLt temperature, heat of fusion, and therefore. 5 crystallinity. This scan also dernonstrates the difficulty in determininq qlass transition tempera ture.
I t i s suspected that the cause of the inaccurate values in Table 4 may have been
created by the chambers of the DSC machine not being cooled down completely in
between the runs of different samples. The heat exposed to the samples when
they were first placed in the pans, prior to beginning the scan, may have been
responsible for pre-melting the samples, thereby changing the glas transition
temperatures and the characteristics of the melt endotherms.
However, the values obtained for the three parent polyurethanes match relatively
closely with data previousty collected by Skarja[74]. Therefore, it i s believed that
the general trends seen are worth discussing, taking into account that there i s a
possible degree of error in the actual values. The only value that does differ
significantly from Skarja's data, i s that of the % crystallinity seen in the
PhelPCL2000. Skarja saw approximately twice as much soft segment crystallinity
for this material.
The pure PEO polyurethanes did not show a melt temperature because they are
completely amorphous, and therefore do not re-order into crystalline structures
when they are heated. They did, however, demonstrate a second transition
similar to a % las transition around 45OC for the PEOlOOO and 55°C for the PE0600
material (see Figure 18). Similar secondary transitions have been seen before by
other researchers, and have been attributed to factors such as the disruption of
ordered, non-crystaliine hard segments, or disruption of interactions (i.e.
hydrogen bonding) between the hard and soft segment[75]. Alternatively, it may
represent a rnixed phase glass transition temperature. Previous work by Skarja
and Woodhouse did not show this phenornenon for the PEO-based polyurethanes,
but it was seen with sorne of the lower molecular weight PCL materials. As it has
now been found in conjunction with PEO as well, it i s likely that this second
transition represents either a mixed phase transition temperature or the disruption
of ordered hard segments, which could be formed for both PEO and PCL-based
materials[76]. These second transitions may be present in the blends as well, but
due to the low resolution, and difficulty pickins up the larger slope change
associated with the soft segment Tg, these secondary transitions may well have
been missed when the cornputer analysis of the spectra was performed.
PEOIOOO
Figure 18: DSC of Phe/PE01000, indicating a smaller transition following the glass transition. suspected to be either a mixed phase transition, or a disruption of ordered hard se2ments.
Al1 of the blends demonstrated melt temperatures lower than that of pure
Phe/PCL2000. Therefore, the incorporation of an amorphous PEO-based
polyurethane appears to lower the glass- transition temperature. The blendinp of
an amorphous polyurethane with a semi-crystalline polyurethane likely disrupted
some of the phase mixing that occurs in the pure semi-crystalline material. Such
increases in phase segregation are known to lead to lower g las transition
temperatures, as was seen here[31].
Normally, an increase in soft segment molecular weight also corresponds directly
with a decrease in glass transition temperature[31]. This trend was s t i l l seen
when the molecular weight of one of the two soft segments in a polyurethane
blend i s increased. In both cases, the increase in molecular weight resulted in a
glass transition temperature decrease of approximately 2 times.
A second trend was seen related to the ~ l a s s transition ternperatures: the higher
the amount of PhelPEO in the blend, the lower the glass transition temperature.
Both of these trends lead to a decrease in glass transition, which i s related to the
amount of mobility available to chains within a polymer. The lower Tg's correlate
to more mobile polymers. Thus, there are fewer factors restricting movernent in
the higher molecular weight and higher %PEO potyurethanes. This can be
attributed to changes in phase mixing, hydrogen bonding, and
hydrophobie/ hydrophilic interactions.
The crystallinity values of the soft segments were relatively low, varying from
approxirnately 24 to 17 %. The PEO-containing polyurethanes are known to be
amorphous, as i s shown by no melt endotherm on the DSCs. Therefore, adding
Phe/PEOs to the PhelPCLZOOO i s expected to induce a decrease in crystallinity of
the soft segment, as was seen. In addition, increasing the amount of PEO content
would decrease the crystallinity even further. This trend was also seen.
lncreasing the rnolecuLar weight of the PEO also served to decrease the %
crystallinity. This i s logical as the PEO chain wi l l be longer, and therefore the PEO
regions wil l be bigger, thereby increasing the volume occupied by amorphous
regions within the material.
3.2.3 Polarized Liqht Microsco~v
By observing the degradable polyurethanes under crossed polarized filters, the
morphologies of the materials were compared. Figure 19 shows Light micrographs
of blends 2 and 4. From these images, it can be seen that the 75/25 blend
exhibits maltese cross patterns. This pattern represents highly ordered crystalline
regions, forming the crosses, with amorphous regions embedded between the
crystal lamellae(32, 771. Blend 4 shows a completely different structure. lacking
spherulites all together. 60th images are very uniform, however, indicating that
the polyurethane blends are weN mixed. It was unexpected that a 25% difference
in PEO-containing polyurethane would lead to a change from a material that i s
highly crystalline to one that i s predominantly amorphous. These morphologies do
agree with the uniaxial tensile testing results, which showed that blend 4 had a
lower rupture stress than blend 2. Since crystallinity imparts strength, the more
arnorphous blends should have lower ultimate tensile stresses. In addition, the
DSC results confirmed that there was a Little more crystallinity in blend 2 than in
blend 4. The differences in c rys tah i ty seen in DSC were not expected to create
such a difference in morphology.
Polarized light micrographs were taken of al1 four blends and the PhelPCL2000
(Figure 20) at a later date. The results seen this second time were significantly
different from the original images. In the more recent work, holes could be seen
in several of the images, particularly those of blends 2 through 4. In the image of
Phe/PCL2000, whitish regions can be seen, bordered by darker edges, which may
represent spherulitic crystalline regions. The typical, spherulitic structure
expected based on the images in Figure 19 of the 75%: 25% blend i s not clearly
seen, but rnay be present. Focussing on the surfaces of the films was difficult due
to the thickness of the films. Layers beneath the surface interfered with obtaining,
a clearer image.
Figure 19: magnification: Phe/PCUWO:
Polarized light microscopy images of desradable polyurethane a) blend 2 (75% Phe/PCL2000: 25% Phe/PE01000), and b)
50% Phe/PE01000).
films blend
Figure 20: Polarized light micrographs of: a) PheIPCL2000, b) blend 1 (75% Phe/PCL2000: 25% PhelPEObOO), c) biend 2 (75% PhelPCL2000: 25% Phe/PE01000), d) btend 3 (50% Phe/PCL2000: 500 PhelPE0600) and e) blend 4 (50% Phe/PCL2000: 50% PhelPE01000). White bars represent 50um.
The samples used for these images were films similar to those used for the
degradation studies, which had been stored in desiccators for at least 3 months.
The images originally taken of blends 2 and 4, however, were taken Less than a
week after casting. It i s therefore likely that the holes that appeared in these
images were not apparent in the first set since the blends had not had as long, to
absorb moisture from the air during storage. The hydrophilic PEO likely absorbed
water into the blends, thereby initiating hydrolytic degradation of the
polyurethane films.
3.2.4 Mechanical Tensile Testinq
Results from the uniaxial tensile tests are summarized in Figures 21 through 23. In
addition, representative stress-strain curves for the blends are shown in Figure 24.
From these results, it i s evident that these materials possess a wide range of
tensile properties.
Figure 21 indicates that Phe/PCL2000 and blends 1 and 2 possess ultimate stress
values closer to those of commercially available wound dressings. The blends
containing the higher PEO contents demonstrate significantly lower rupture
strains. This i s likely due to the fact that PEO-containin9 polyurethanes are rnostly
amorphous, whereas PCL-based materials possess a higher degree of crystallinity.
By adding more PEO-based polyurethane to the blend, the crystalline regions will
likely be smaller. The strength of a polyrneric materiat i s greatly affected by i t s
degree of crystallinity. Since crystalline structures have an ordered structure,
interactions between molecules, such as hydrogen bonding, can be maxirnized,
thereby imparting greater strength. Therefore, these resul t s demonstrated in
Figure 21 are not unexpected. It i s also important to note that while blends 3 and
4 have lower ultimate stress vahes than commercially available dressings, they
s t i l l possess enough strength to act as a wound dressing. A wound dressing should
be at least as strone, as skin, which has an ukirnate tensile stress of 15 MPa[78,
791. Therefore, al1 of the materials tested possess more than enough strength for
the proposed application.
It is a Little surprising that for the 75/25 blends, the PE01000-containing blend was
stronger, but in the 50/50 blends, the PE0600 blend was higher. This trend i s seen
in the data for ultimate tensile strain and Young's moduhs as well, though the
difference i s not significant for strain. In making the blends, as the polyurethanes
were weighed out prior to filtering, it i s possible that the final ratios of materials
in each blend may differ from the expected. These discrepancies could very well
account for the differing trends.
ALI five of the degradable polyurethanes dernonstrated comparable or greater
percent elongations a t rupture. Thus, any of these materials would provide
enough stretch to allow for movement of joints and easy application. The
degradable materials can be stretched to over 500%. The stress-strain curves in
Figure 22 show, however, that only approximately 5-15% elongation can be
achieved before the stretching becomes irreversible for the blends. This result
was also seen for PhelPCLZOOO (not shown). This could mean that i f the dressing
were stretched too much over a joint, it rnight not return to i t s original size,
leading to possible unwanted excess material over the wound.
Al1 of the degradable polyurethanes exhibited Young's modulus values (Figure 23)
of at least four times those of both 0psite3 and ~egaderm'. The Young's modulus
of human skin i s 35 MPa, a value higher than those of the commercially available
wound dressings, but lower than al1 five polyurethanes[78]. This indicates that
more force i s required to reversibly elonsate the degradable polyurethanes a siven
distance than either skin or opsite' and ~egaderm? Since this i s the case, it is
expected that a dressing made of any of the blends i s unlikely to get stretched so
rnuch that it wi l l not return to i t s original state.
The stress-strain curves depicted in Figure 24 al1 have very similar curve shapes.
The blends containing more PEO exhibit curves that are lower and shifted slightly
to the right. They are also unable to be stretched as far, as discussed above. This
indicates that the 50150 blends can be stretched farther for a given force than the
corresponding 75/25 blends, but wil l break at shorter elongations.
Ultimate Tensile Stress of Polyurethane BJends and Controls
Blend4 Y i +
Figure 21: Ultirnate tensile stress values for the polyurethane blends, as well as pure Phel PCL2000 and two commercially available polyurethane dressings. Each bar represents an average of 6 values. Error bars represent 95% confidence intervals.
Ultimate Tensile Strain of Polyurethane Blends and Controls
Strain (% elongation)
Fisure 22: Ultimate tensile strain vatues for the polyurethane blends, PheIPCL2000, 0psite"nd ~egaderm". For each material, n=6. Error bars represent 95% confidence intervals.
Young's MDdulus of Polyurethane Blends and Controls
Young's Modulus (Mla)
Figure 23: Young's Modulus values for the polyurethane blends. Phe/PCL2000. 0psiteg and ~egaderm '. For each ma terial. II-6. Error bars represent 95"$onfidence intervak
Figure 24: Typical tensile stress-strain curves for the four poiyurethane blends.
3.2.5 Deqradation Results
Total mass loss and appearance under the scanning electron microscope were used
to determine the effects of buffer degradation at 37'C on the polyurethanes. A
summary of the percent mass loss results for the first four weeks are shown in
Figure 25. Measurements were also taken after 6 months, but the mass loss values
were not significantly different from the 4-week values for all four blends (95%
confidence interval, student's T-test). The PhelPCLZOOO samples for 6 months
showed a drastic decrease in total mass loss (to 0.7 9 6 ) . If the 6 month samples
were identical to those used for the other data points, there is no viable
explanation for this decrease. Perhaps a different cast film was used when
preparing samples for this time point. A repeat of the study would provide more
insight in to this issue. However, an enzymatic degradation study would prove
more useful in determining actual mass loss rates under physiological conditions.
There was a high degree of variability in mass loss for blends 2 throush 4. The
Phe/PCL2000 and blend 1 appeared to (ose most of their mass in the first seven
days, and then levelled off. The other three blends showed an overall trend of
levelling off foilowing one week as well, though these results were more varied.
Logically, mass loss should not be able to decrease with tirne. I t i s suspected that
the small initial mass values have led to a significant source of error. On average,
samples started off weighing approximate 3-6 mg, depending on the material and
thickness of the film. For a 3 mg sample, a 5% mass loss represents a mass change
of only 0.15 mg. The analytical balance used in these studies was only accurate to
one decimal place in mg mode. Therefore, the dips seen in the mass loss curves
may be due simply to imprecise mass values. For instance, degraded pieces of the
polyurethanes may have adhered to the surface of the discs, thereby skewing the
actual mass reading of the sample. Repetition of this study, using larger sampLes
could be used as a method of reducino, the sources of error.
In addition, even though the samples are dried for 7 days prior to weighing, the
polyurethanes may have absorbed some additional water which was not removed
during this time. This could make the mass loss smaller than it really is, or in
some extreme cases, could even make the mass loss appear to be negative. Thus,
the variations in mass may very well be due to differences in hydration of the
various sam ples.
Based on the results from this study, it is evident that al1 of the blends degrade
quite slowly, on the order of months to years. In vivo, however, these rates could
be sped up through enzymatic degradation. Since the percent mass loss values
appear to level off more or les , it i s likely that the PEO-containing polyurethane
component of the blend i s hydrolysed rapidly, leaving behind the much more
stable PCL-based material to degrade significantly slower.
Mass Loss as a function of Time - LDVPhe/PCL2000 -- -
Blend 1
5 1 O 15 2 O 25 30 Tim e (days)
Figure 25: Mass loss as a function of tirne for the degradable polyurethane blends m=41 and Phe/PCL2000 (n=4). Each data point represents average total mass loss since day 0.
Figure 26 through Figure 30 show representative scanning electron micrographs, at
lOOOx masnification, for each material, indicating the appearance of the materials
prior to degradation, after 2 weeks, and after 6 months. From these figures it i s
evident that white there i s no visible degradation of PhelPCL2000, the blends al1
under50 deqadation to some extent.
In Figure 26, there i s only visible degradation at the edges of spherulitic regions in
Phel PCL2000. These regions, seen as approximately hexagonally shaped regions,
appear to be a l i t t le vulnerable to degradation at the edges, but do not show any
surface degradation elsewhere. The presence of spherulitic regions in these SEMs
indicates that perhaps spherulites were what was visible using polarized light
microscopy as well. This would agree with the first set of polarized micrographs
that demonstrated spherulites in the 75% PhelPCL2000 blend. Since the PEO-
containing polyurethane i s amorphous, any crystallinity imparted to that blend had
to have been a result of the PhelPCL2000.
Al1 four of the blends (Figure 27a - Figure 30a) demonstrate an interestins and
unexpected phenomenon: even before degadation begins, the films appear to be
permeated with voids. These holes seem extremely uniformly distributed, and
very circular in shape. In Figure 27a, it appears that the voids in blend 1 are
actually crevices rather than complete holes. In blends 2 throuph 4, both crevices
and holes can be seen. The size of these voids varies for the different blends. For
blends 1 and 2 , they are on the order of 2 pm in diarneter on the surface. Blends 3
and 4 contain holes approximately half that site. Thus, the higher PEO content
polyurethanes contain smaller voids. They appear to contain at least twice as
many holes per unit area, however, indicatins that though the voids are smaller,
the void space may be higher. There are two likely causes of these holes. The
first i s the possibility that water vapour was pulled out of the air and into the
polyurethane-chloroform mixture either prior to or during casting. The second
possibitity is that after casting, the hydrophilic nature of the FE0 within the
blends may have attracted and adsorbed water into the film, beginning the
degradation process. Al1 of the samples were kept in a dessicator, but could have
been exposed to some humidity, especially during the summer months. 80th of
these explanations focus on water being attracted to the polyurethane, causing
premature degradation. It i s very possible that both of these mechanisms played a
role in the formation of voids within the blends.
It i s evident from Figure 27 that as degradation proceeds, more of the crevices
open up to form holes in the polyurethane. As degradation time increases. the
amount of solid polyurethane around these areas decreases. No visible
degradation i s apparent in these solid areas surrounding the voids. Blend 1 starts
with a larger ratio of solid polyurethane regions to total area of the material than
blend 2. After 2 weeks and 6 months of degradation, however, the reverse i s true.
Blend 2 demonstrates a highly unexpected degradation trend (Figure 28): the size
of the voids and holes decreases with time. The distribution of these features
appears to remain uniform, however. Likely, what i s happening in this case i s that
when exposed to a hydrated environment, the PEO-containing sections may
reposition themselves on the surface of the film, hiding the Less hydrophilic
portions of the polyurethane within. After 2 weeks of degradation, most of the
PEO may have been hydrolysed and removed, leaving behind a rnaterial much
richer in PCL. The few srnall voids seen may indicate where small resions of PEO-
rich polyurethane are ernbedded within the PCL-containing elastomer. These
results would agree with those obtained by Skarja and Woodhouse[71], which
showed that the PEO-based materials were more susceptible to degradation and
erosion than the PCL-based materials. This trend was seen even though the PEO
soft segment i s not as highly susceptible to hydrolysis as the PCL[71]. It was
hypothesized that the reorientation of the PEO polyurethane surfaces, and the
hydrophilic nature of the PEO within it, was responsible for recruitina more water,
and thereby improving the hydrolytic degradation rates.
As seen in Figure 26, relatively l i t t le degradation affects the PCL poLyurethane
surface, except at the edges of spherulitic regions. This same trend of Little
noticeable degradation i s seen for blend 2 between 2 weeks and 6 months. The
ridges seen in Figure 28c are evidence of marks transferred to the film by the
casting dish used in casting the film. Various sizes of casting dishes were used,
and the largest dish was machine polished, leaving grooves in the Teflon that
appear to affect the surface roughness of the films obtained. On samples that
showed these grooves, fewer voids were seen, indicating that surface roughness
may affect and prevent degradation. There are no grooves seen on the O day and
2 week samples, since they were taken from a different film. The only difference
between the O day and 2 week time points appears to be a decrease in the size of
voids, indicating a further removal of the PEO polyurethane from the surface.
Blend 3 (Figure 29) also shows an interesting degradation pattern. A t day 0, the
appearance i s similar to the other blends. As degradation proceeds, however, it
appears that two or more voids become joined, with smaller voids formed beneath
them. The appearance of the surfaces after 2 weeks and 6 months is very similar.
indicating that the surface rearrangement that may occur in blend 2 does not
happen with blend 3. This trend continues for blend 4 (Figure 30), although the
voids do not appear to join as they do for blend 3. They do increase in size from
day 0, however, indicating that degradation i s occurring.
Figure 26: SEMs of Phe/PCU000 at 1000x magnification: a) before degradation, b) after 2 weeks exposure to buffer at 37'C, and c) after 6 months exposure. Btack bars represent 30~tm.
These SEM imases demonstrate that surface degradation of the blends appears to
centre on the formation of voids within the material, likely an effect correlated
with PEO content within the polyurethanes. Thus, the mass loss data levelling off
66
does indeed appear to correspond to the removal and Levelling off of PEO content
within the material. In future, analysis of the mechanical properties and
determination of the composition of the materials post-degradation might yield
useful information as to how much PEO-rich material i s lost as a result of
degradation.
Figure 27: SEMs of blend 1 (75% PhelPCLZ000: 25% Phe/PE0600) at 1000x magnification after: a) O days, b) 2 weeks, and c) 6 months degradation. Bars represent 30prn.
Figure 28: SEM images of blend 2 (75%PhelPCLZ000: 2516 Phe/PE01 OOO), at 1 OOOx magnification. following: a) O days, b) 2 weeks, and c) 6 months degradation times. Bars represent 30 um.
Figure 29: a) O days, b) 2 weeks, and c) 6 months degradation of blend 3 (50% PhelPCL2000: 50% PhelPEObOO), as seen using SEM, 1000x magnification. Bars represent 30 ~ im .
Figure 30: SEM images of blend 4 (50% Phe/PCt2000: 5û% Phe/PE01000), following a ) O days, b) 2 weeks, c) 6 months degradation, IOOOx magnification. Bars represent 30 urn.
Chapter Four:
Development of the Fibrin Coating
4. O Introduction
Fibrin sealants have been used to help promote healinz and enhance dressing
adherence in cutaneous wounds for decades[26, 39, BO]. As cells begin to enter a
wound site, a fibrin scaffold must first be laid down to allow for and guide cell
movement[4]. As a result, coatins a wound dressinp with a thin fibrin film could
provide a means of allowing for faster infiltration of cells into the wound site. A
fibrin coating method, originally developed by Rubens e t al. for blood-contacting
applications[81-831, was recently adapted for wound dressing applications[37, 841.
The coatine, protocol involves the formation of a two-layer coating (Figure 31).
The bottom layer, composed of thermally denatured fibrinogen, provides a method
of anchoring the coating to the polyurethane substrate. The second layer i s
composed of fibrin, which has been cross-linked for added stability.
Cross-linked fibrin
1 j- Polyurethane substrate
figure 31 : The fibrin coating, developed by Rubens et a1.[83]. Two distinct layers are coated onto the substrate: first, a thermally denatured fibrinogen layer (TDF), fotlowed by a layer of polymerized, cross-linked fibrin.
The substrate materials used in previous development of the fibrin coating were
~egaderm", Corethane@ (a woven polyurethane used in vascular grafts), and one
of Skarja's earlier degradable polyurethanes[30] that did not contain a lysine-
based diisocyanate[37, 841. Bense's work showed that thin, uniform fibrin
coatinçjs could be produced on both flat and woven polyurethane surfaces. Based
on these results, it was hypothesized that a uniform coating could be applied to
the newly developed degradable elastomers of Skarja and Woodhouse[34].
4.1.1 Materials
The materials used in development of the fibrin coatinps are listed in Table 5. In
addition to these materials, iodine-125 radiolabelled fibrinogen prepared by Glenn
McCLung at McMaster University, Hamilton, ON, Canada was used. Prior to use, the
fibrinogen received from Calbiochem was dissolved and dialysed azainst tris
buffered saline (TBS), and i t s concentration was determined using, ultraviolet
spectrophotometry.
Item Source
D-Phe-Pro-Arg Calbiochern, La Jolla, CA Ketone (PPACK) , dihvdrochloride , , -- .
Factor XII 1, human Enzyme Research Laboratories, South Bend, IN
Fibrinogen, Oregon Green 488 Molecular Probes, Eugene, OR la belled Fibrinogen, plasminogen depleted, Calbiochem, La Jolla, CA from human plasma
. -
Throm bin Sigma, St . Louis, MO
Table 5: Sources of materials used in creation of the fibrin coatins.
4.1.2 Fibrin Coatina Develo~ment
The fibrin coating protocol was adapted from that of Bense and Woodhouse[M].
Bense suggested that cross-linking times should be investigated, and possibly
extended in order to improve the stability of the coatins[37]. Initial work on
adaptine, the coating was performed usin3 ~egaderm" a material already used by
Bense, in order to allow for comparisons to be made. Changes were made to the
protocol as needed, based on observations made durina the coating procedures,
and known shortcomings seen in previous work[37, 84, 851. The final coating
protocol developed i s illustrated in Figure 32, below.
Punch circular disks, rinse in methanol a -- - 7.1,7,,~2 - - ; L m - - . J 2
Soak in TBS for 1 hour
Place in 1 mg/mL fibrinogen in TBS for 2 hours
v' Incubate a t 70°C for 15 minutes
Soak in 1 pg/mL thrombin in hypotonic TBS for 5 minutes
a Rinse 3 times in hypotonic TBS
S o a ~ in C . 5 in(; r r L ribr:nqe!? 111 T i jS 'or ? 5 mrnl;;e:
a Rinse 3 times in hypotonic TBS
u Rinse 3 times in hypotonic TBS
Soak in 10 urnol/L PPACK in TBS for 5 minutes
Figure 32: The new fibrin coating protocol. Steps described in purple text represent changes made to the previous protocols developed by Rubens et al. and Bense and Woodhouse[37. 831.
In order to determine the reproducibility of the coatings, and to ensure that fibrin
was being coated onto the surfaces, radiolabelled fibrinogen was used in the 0.5
mg/mL fibrinogen step. Roughly 20% of the fibrinogen in the solution was
radiolabelled, and the remainder was unlabelled fibrinogen.
For assessrnent of uniformity of the fibrin coating, Oregon green labelled
fibrinogen was used instead of radiolabelled. In these cases, 0.5% of the
fibrinogen was fluorescently labelled. Following coating, the surfaces were then
fixed in 10% formalyn solution (Sigma, MO) overnight, and then mounted on glass
microscope slides.
4.1.3 Characterization of the Coatinq Presence and Reproducibility of the Cootings:
Samples were prepared using radiolabelled fibrinogen, as described above, and
analysed using a gamma counter. Background counts, as determined through the
use of controls, were subtracted from al1 values.
Fluorescence Microscopy:
The same microscope and software used in observing the polyurethane blends
under polarized light filters (see section 3.1.3) was used, with a fluorescent light
source.
Scanning Eiectron Microscopy:
The same SEM equipment, discussed in section 3.1.3, was used to observe the
coated samples. Prior to microscopic evaluation, however, the samples had to be
prepared. The coated surfaces were dehydrated through a series of graded
ethanol rinses. Essentially, the samples were soaked in ethanol solutions for 10
minutes each, starting a t 30% ethanol, and ending with 100% ethanol. A total of 6
steps were involved in the process. Once dehydrated, the coated polyurethanes
were mounted and prepared following the same methods described in 3.1 .3 .
4.2 Results and Discussion
4.2.1 Coatinq O~timization Over the course of this study, several modifications to the fibrin coating, protocol
were made. Initially, a Longer cross-linking time of 1 hour was chosen based on
previous results[85], and the cross-linking temperature was changed to 3 7 O C in
order to op timize the enzyme's activity . A t p hysiological tempera tures, a solution
containing both fibrin and factor Xlll results in a highly viscous solution.
corresponding to the formation of a cross-linked hydrated clot. Removal of coated
samples from this solution was difficult, resultinp in a gel-like capsule around the
surfaces. Either this entire capsule was carried through the remainder of the
coating protocol, resulting in a very thick coating that i s quite easy to remove, or
the coating was removed in an attempt to remove the excess fibrin gel around the
surfaces. Based on these results, the formation of the cross-linked fibrin layer was
converted into a two-step process. Here, the fibrinogen i s first adsorbed to a
thrombin rich surface, and thereby converted to fibrin. Next, the surfaces are
incubated in a factor X l l l solution, with twice the concentration of FXlll used by
Bense and additional thrombin, to ensure that the factor Xlll cross-linking reaction
i s encouraged. These chanpes were originally performed on ~egaderm', and the
reproducibility of the coatings was ensured utilising '251-(abelled fibrinogen. For
these studies, eight samples were coated and the number of counts per surface
were compared. Numerous attempts were made to create a calibration curve so
that these cpm (count per minute) values could be converted into mg/cm2.
However, a reproducible and accurate calibration curve could not be obtained due
to fluctuations in the sensitivity of the gamma counter. Therefore, al1 values for
these investigations were left in the units of counts rather than mg of fibrin(ogen).
The values within each batch on ~ e ~ a d e r m ' indicated acceptable reproducibility . Once these coatings were successfully completed, attempts were made to coat
PhelPCLZOOO and the first two blends created. The reproducibility on these
materials was found to be very poor. The number of counts per surface were seen
to Vary by more than 2 orders of magnitude. An example of typical data obtained
i s illustrated in Table 6, below. Since different batches of fibrinogen were used
for the different materials, cornparisons of actual counts per minute should not be
made between material types. Only variance within each batch could be
compared .
Material Counts per Minute (mean)
Tegaderm l 3657 2 11 04
Phe/ PCL2000 1
I 347161 i 625873 l
Blend 1 13421 i 74533
Blend 3 7379 - 13073 1
Table 6: Sample ranges of count per minute values obtained for surfaces coated with the fibrin coating using the original protocol. For each material, n=9. The values represent the means * 95% confidence intervals.
In order to determine where the coatine, was becoming inconsistent, the protocol
was broken down and, usin5 1251-labelled fibrinogen, the samples were coated with
only the adsorbed fibrinogen layer, prior to denaturation, and counted. For these
studies, Phe/PCLZ000 and Blend 1 were utilised. The results indicated that the
irreproducibility of the coatings was occurring with the placement of the TDF
Layer. I t was hypothesized that residual methanol remaining on the surfaces might
prevent uniform adsorption of the fibrinogen. A series of protocok, whereby the
surfaces were rinsed three times in TBS or Triton X-100 prior to the first buffer
equilibration step, was developed. In using '251-labelled fibrinogen for the first
fibrinogen layer, much better results were seen when the materials were rinsed in
TBS prior to coating.
Step after which countinq was Counts per minute (mean)
performed
Triton X-100 rinse, first fibrinogen Layer '
9274 4 6972 absorbed, not yet denatured
Methanol, first fibrinogen layer 15056 2 81 94
absorbed, not denatured
Methanol, TBS rinse, first fibrinogen 15965 2 3138
Layer absorbed, not yet denatured I
Table 7: Effect of different rinse protocols on PheIPCU000 radiolabelled fibrinogen adsorption. For this study, n=9. Values represent means k 95% confidence interval.
Step after which countinp was 1 I Counts per minute (mean) performed
Triton X-100 rinse, first fi brinogen layer 86231 5841 6
absorbed, not yet denatured
Methanol, TES rinse, first fibrinogen layer 3672 i 983
absorbed, not yet denatured
Table 8: Effect of different rinse protocols on BLend 1 (75% Phe/PCL2000 : 25% Phe/PE0600) followinq '25~-fibrinogen adsorption. n-9. Values are expressed in means 95"6onfidence interval, as determined using student's standard t-test.
Once the protocol had been improved, the coatings were then carried out as
before, with 1251-labelled fibrinogen in the fibrin layer rather than the TDF layer.
and again reproducibility was better. Values s t i l l varied significantly more than on
~egaderm', but were much better than they had been, and varied by Less than one
order of magnitude. Finally, al1 five materials were coated using the new method
to determine reproducibility of the final coatings (Table 9). These new coatinss
s t i l l demonstrated a fair amount of variability between surfaces. The variability is
particularly poor for the 50:50 blends. This may indicate that hydrophilicity or
morphology of the blends will affect how reproducibly the materials can be
coa ted.
Materiat Counts per minute (mean)
P he 1 PCL2000 7341 26 5 291 27
Blend 1 47588 I 6203
Blend 2 773393 I 63774
Blend 3 4268 I 31 28
Blend 4 98492 + 36214
Table 9: Final counts per minute of various degradable polyurethanes coated with radiolabelled fibrinogen. Values represent mean t 95% confidence interval.
4.2.2 Fluorescent Liqht Microscoo~ Al1 four blends, ~egaderm', and PhelPCLZOOO were coated using Oregon Green
labelled fibrinogen. When observed under the microscope, al1 surfaces exhibited
fluorescence throughout the surface, indicating that the entire surface had been
coated with fluorescently labelled protein. Some areas fluoresced more brightly,
indicating a higher concentration of fibrin in those locations. These results were
seen for al1 six types of surfaces.
Figure 33 dernonstrates a uniformly coated surface, and one where some denser
fibrin patches can be seen. The images became blurred at rnagnifications of more
than 10x. For some samples, more than 5x was not focusable. Regardless of the
magnification, the fluorescent content was high enoush to inhibit resolution of the
fibrin coatings. In an attempt to minimize this effect, different amounts of
labelled fibrinogen were incorpora ted into the coatings. Even at the smallest
accurately measurable quantity (O.OSmg), the coatings demonstra ted the same
appearance. Thus, beyond determining, that the entire disc became coated with
fluorescent protein, very l i t t le information as to the nature of the fibrin could be
discerned usinp this method.
Figure 33: Fluorescent h ~ h t micrographs of Oregon Green Labelled fibrin coatings. A) Blend 1, as seen at 5x magnification. B) Phe/PCL2000, at 10x masnification, indicating more fluorescent fibn'n in some places, creatins a wispy pattern.
4.2.3 Scannina E l e H
When the coatings were observed using SEM, varyin5 degrees of coverage were
seen. On several samples, the coating was patchy, Leaving large segments of
polyurethane exposed. This was seen on al\ six material substrates, and may be an
artefact of the sample preparation methods used. Figure 34, below, demonstrates
the patchy nature of the coating. Where the coating i s present, however, it does
appear to be quite weU anchored, as is dernonstrated at the edses of Figure 35.
Figure 34: SEM of fibrin-coated blend 1 (75% Phe/PCL2000: 25% Phe/PE0600), a t 25 times rnagnification. The lighter areas represent the fibrin coating, whiIe the darker regions are the underlying polyurethane substrate. Bar indicates 1.2 mm.
The fibrin coatinps dernonstrate a fibrous network characteristic of fibrin clots at
high magnification (Figure 36). On the blends, the coating appears to follow the
contours of the depressions and holes within the surface, seen as divots in the
coating. At 5000x magnification, the fibrin coating on blend 3 appears quite
patchy. The fibrin appears to have formed within the voids of the polyurethane
substrate as well as on the surface. The coating is not smooth and uniform in rnost
places, as seen in the images for the other five materials. However, the above
SEM (Figure 30) demonstrates that on some portions of the blend 3 surfaces, a
cohesive coating i s formed.
Figure 35: SEM at 1 OOOx magnification, showing an area where the polyurethane has become bent, creatins a Sap between the coating and material. The substrate material seen in this figure is blend 3.
In taking the SEMs, attempts were made to only capture images that were
representative of features seen on al1 samples, rather than anomalies. Thus, the
other five materials appear to support a more uniform, cohesive fibrin coating
than blend 3. In developing a fibrin coated wound dressing, therefore, one of the
other blends should be used so as to allow for a greater amount of fibrin to be
present on the surface of the biomaterial. Blend 3 miaht, however, be well suited
to the fabrication of fibrin coated porous three-dimensional scaffolds for other
biomedical applications. See Appendix Three for a more detailed description of
three-dimensional foam formation using the blends.
Figure 36: SEM'S of fibrin coated polyurethanes, 5000x magnification: A ) ~egaderrn" 0) PhelPCL2000, C) BIend 1 , D) Blend 2, E) Blend 3, and F) Blend 4. Bars represent 6 ~ im.
On surfaces that possessed machining grooves that were transferred from the
casting dish, the fibrin coating foUowed these contours, and appeared to build up
thicker coatings along these grooves. This behaviour is demonstrated in Figure 37,
below. These results are encouraging, as it indicates that during coating, the
fibrin appears to find and migrate into regions that are elevated as well as into
depressions, as discussed earlier.
Figure 37: Fibrin coated blend 2, seen a t 500x magnification. The polyurethane film contains grooves upon which the fibnn coating has build up.
Chapter Five:
Incorporation of Hyaluronan into the Coatings
5.0 Introduction
Hyaluronan (HA) affects several aspects of wound healing. Its important role in
prevention of scar tissue formation i s of particular interest to cutaneous wound
applications. Not only function, but also cosrnetic appearance, i s an issue for
these wounds. Through the use of exogenous HA on wounds, researchers have
been able to show improved angiogenesis, less wound contraction, a direct factor
in scar formation, and increased cellular infiltration.
In addition to hyaluronan's attractive properties, it also binds specifically to fibrin
and fibrinogen. Therefore the addition of HA to the coating i s expected to not
only alter the wound healing response, but also, may affect the nature of the
coating.
This chapter of the thesis deals with the incorporation of hyaluronan into the
coatings developed in Chapter Four, and the subsequent characterization of these
new coatings. As with the fibrin coatings, in creating the coatings, the protocol
previousty used by Bense[84] needed to be adapted.
5.1 Experimtal Methods
S u l Materials The materials used in developing the fibrinlHA coatings are the same as those
used for the fibrin coatings. In addition to these materials, pharmaceutical grade
hyaluronan was obtained from Hyal Pharmaceutical Corporation, Toronto. This HA
had a molecular weight range of roughly 500-800 kDa.
5.1.2 Fibrin/Hvaluronan Coatinq Deveîo~ment The coating protocol used is the same as that described for the fibrin coatings,
with the addition of 6 pg/mL HA to the 0.5 mg/mL fibrinogen adsorption step.
Near the start of this research, two different protocols were investigated: one
that added the fibrinogen and HA at the same time, and another where the
fibrinogen was allowed to adsorb for 15 minutes, and then the hyaluronan was
added for an additional 15 minutes.
5.1.3 Characteriration of the Coatinqs In characterising the coatings, light microscopy, and SEM were utilised. As with
the regular fibrin coatings, the earlier SEMs, performed on Tegaderm were done
using a freeze-drying technique to dehydrate the samples rather than graded
ethanol washes. Fluorescently labelled fibrinogen was used in the same
concentrations as for the plain fibrin coatings for determinine, the effect of
hyaluronan on the distribution of fibrin within the coating. A cationic stain, Alcian
blue, was also used to identify the distribution of HA in the coatine, using Light
microscopy. In performing these studies, the samples were coated using
unlabelled components, and then were incubated in a 0.05 % (w/v) solution of
alcian blue in 2 % ( v l v ) acetic acid for 5 minutes. Following this stainine, period,
the samples were destained in 2% acetic acid for times varyinp from 5 minutes to
48 hours.
5.2 Resulis and Discussion
5.2.1 Scannina Electron Microscopv As discussed in section 5.1.2, original HA-con taining coatings were produced using
two different methods. Both coatings on ~egaderrn" were observed using SEM.
The two resultant coatings are shown in Figure 38. These imases show two
coatings that look very similar. However, the coatings made with the HA added
after the fibrinogen had adsorbed for some time tended to be more patchy, with
Large areas of the surfaces uncoated. This may be a result of the hyaluronan
remaining in solution and binding to the fibrin coated ont0 the surfaces. The
hydrated HA network may have pulled the fibrin back off the surface. In addition
to these observations, results reported by Nehls et al. indicated that hyaluronan
must be present during fibrin formation rather than after in order for improved
cell migration to be seen[66]. Based on these two factors, the protocol in which
the fibrinogen and hyaluronan were added at the same time was adopted for al1
further studies.
Figure 38: FibrinlHA coatings on Tegaderm performed using the methods of A) adding the fibrinogen and HA at the same time, and 0) adding the fibrinogen first. Magnifications are 1150x and 1000x, respectively. The ovoid shaped particles are contaminants (dust). Bars represent 30 L m .
Al1 four degradable blends and PhelPCLZOOO were coated with fibrin and HA usin5
the selected protocol. The coatings produced are shown in Figure 39. These
coating were much more extensive and uniform than those made with fibrin
alone. These coatings appear to blanket the surface, filling in, and following the
contour of some surface features such as mal t crevices and holes. This
phenornenon i s likely due to the hygroscopic nature of HA. Hyaluronan in the
fibrinopen solution wil i swell in addition to b ind in~ to the hyaluronan , providing a
gel-like matrix that will surround the TDF coated substrate. By swelling and
surroundinq, there
wil l be exposed to
is a
the
the protein molecules
causing inconsistencies
better chance that al1 regions of the polyurethane surfaces
same microenvironment. In the fibrinogen solution alone,
may clump together or stick to the apparatus, thereby
in the coating appearance alon2 the surface.
Figure 39: SEMs of fibrin and hyaluronan- coated biodegradable polyurethanes: A) Phe/PCL2000, 8) blend 1, C) blend 2, O) blend 3, and E) blend 4. Al1 images are 5000x magnification, except for blend 3, which i s shown at 2000x. White bars represent 6 h m .
In Figure 39b), an interface between the coating and blend 1 can be seen. Such
interfaces were seen on al1 samples, though their size and numbers were
noticeably lower for the fibrinlHA coatings. By focussing in on interfaces such as
this, the thickness of the coating and how attached it i s to the surface can be
assessed qualitatively. Figure 40, below shows a fibrin/HA coating and a fibrin
only coating on blend 3. From these images, it can be seen how the fibrin/HA
coating follows the contours of the polyurethane quite closely. The upper right
hand corner of the fibrin only coating demonstrates that only small features are
filled by the coating. The fibrinlHA coatins does appear to be a Little thicker than
the fibrin only coating. However, it is difficult to compare, as one coating i s flat,
while the other i s l i f ted up. I t appears that the coatings Vary from roughly 1 to 2
pm in thickness, a relatively thin value.
Figure 40: SEM images of btend 3 coated with A: fibrin and hyaluronan, and B: fibrin only. The top of both images represents areas that have been coated, while the bottom portion of each image dernonstrates the bare polyurethane.
5.2.2 Liaht Microsco~y Fibrin and hyaluronan coatings were made on al1 six materials using Fluorescently
Labelled fibrinogen, and observed under the microscope. Like the fibrin only
coatings, the resolution was poor at high rnagnifications. The surfaces appeared
to be well coated, however, as fluorescent green was seen covering the entire disk
at lower magnifications(Figure 41 ).
Figure 41: Blend 1 , coated with HA and Fluorescently labelled fibrin, a t 5x magnification.
Coatings made usin2 unlabelled fibrinogen, and stained using alcian blue were also
observed under the microscope. Upon staining, the samples (including uncoated
controls, and fibrin only coated materials) became a very rich blue. However,
destaining resulted in either no removal, or removal of al1 colour from the
samples. Several different destaining times were investigated, but with all these
attempts, the assay was unable to distinpuish between the HA coatings and the
controls. Since the alcian blue dye is cationic, it should attach itself to the
sulphate groups of hyaluronan, and remain attached following destaining. Samples
that do not contain these negatively charsed groups ought not to retain the stain.
Higher concentrations of alcian blue were attempted to see i f the stain was not
working because it was too dilute. However, the same trend was seen, until a
concentration of 0.5% (wlv) was reached, at which point the surfaces retained
some of the btue die, but so did the fibrin coated and uncoated controls. It is
possible that the acidic conditions of the alcian blue solution was causing
degradation of the polyurethanes to take place, thereby creating pockets within
which the die could be retained, making it difficult to qualitatively distinguish
between the two coatings and the controk A more specific assay for hyahronan.
possibly using biotinylation, or some other similar technique, may be necessary to
determine the distribution of hyaluronan on the surface[Bb]. However, since the
SEMs do indicate differences in coating structure, it i s evident that HA i s affecting
the formation of the fibrin coating. Since fibrin and HA bind specifically,
hyaluronan i s most likely present on the surface. Also, since the HA helps to
distribute the coating more evenly, and it is known to affect healin2 in desired
ways, it should continue to be incorporated into the fibrin coating.
Chapter Six:
Conclusions, Contributions and Recommendations
6.0 Summary In this research, four degradable polyurethane blends were produced and
characterized. These materials were then coated with thin Layers of both fibrin
and fibrinlhyaluronan. The results of this work indicated that degradation and
mechanical properties of biodegradable polyurethanes could be manipulated by
creating blends. In addition, these novel polyurethane blends could be coated
with fibrin and fibrinlHA. These coated materials could represent a viable
alternative to currently available wound dressings that will help enhance the
healing process.
Con tributions Four novel biodegradable polyurethane blends were produched. These
materials may have potential uses in numerous different biomedical
applications, particularly those requirine, an elastomeric material.
The effects of blend composition, and therefore morphology and
hydrophilicity , on mechanical and degradative properties were assessed.
Fibrin coatinp protocols were developed and implemented, resulting in a
reproducible method of creatinçj fibrin coated polyurethane films.
Hyaluronan was incorporated into the fibrin coatinps, and its effect on
these fibrin networks was investigated.
6.2 Concfusions 1. The properties of the blends appeared to be related to the morphologies of
the polyurethanes. The relative compositions of the various btends had a
direct impact on the morphologies, and therefore on the mechanical
properties of the materials.
2.
3.
4.
6.3 There
Chanses in the fibrin coating methodology allowed for a more reproducible
coating within batches to be created on the new polyurethane blends.
The fibrin coating was found to be patchy on the degradable polyurethanes.
Where there i s coating, however, it appears to be well anchored to the
underlying material.
Microscopic evaluation of fibrin/hyaluronan coatings demonstrated thicker,
more uniform coatings than the fibrin alone.
Recummendations for Future Investigation are many issues that have become apparent during the course of this
research. Prior to use of these biomaterials as a wound dressing, some or al1 of
the following should be addressed:
1. The blends need to be characterized further. In particular, the actual
composition of the blends should be determined more accurately, and the
effect of degradation on tensile properties should be investigated.
Repetition of the buffer degradation study, using larger polyurethane
specimens could also be done. In addition, in order to determine
degradation rate more accurately, enzyme degradation studies should be
carried out, using trypsin- and chymotrypsin- like enzymes, to see how
much faster the materials degrade in the presence of enzymes.
2. Fvrther characterization of the fibrin and hyaluronan coating should be
performed, in order to more accurately determine the presence and
distribution of hyaluronan within the coating. Quantification of fibrin in the
coatings should also be performed.
3. Work by Bense[84] indicated that when hyaluronan i s incorporated into the
coating, it becomes more susceptible to plasmin-induced fibrinolysis. The
coating protoco( was altered, however, and more extensive cross-linking is
present in the new coatings. Thus, the susceptibility of the new fibrinlHA
coating to plasmin degradation should be assessed in vitro.
4. In vitro cell culture techniques should be utilized to determine the effects
of the bare degradable polyurethanes, and the coated polyurethanes (with
and without fibrin), on cell migration, angiogenesis, and wound contraction.
5. In vivo evaluation of the polyurethane blends and fibrin (Ihyaluronan)
coatings should be performed, as a follow up to the preliminary study
discussed in Appendix Four. A non-rodent healing mode1 should be
considered.
References
Thomas, S. and I .M. Leigh , Wound dressings, in Wounds: Biology and Manaqement, D.J. Leaper and K.G. Harding, Editors. 1998, Oxford University Press, Inc.: New York, p. 166-183. Holbrook, K.A., Structure and Development of the Skin, in Pathophysiology of dermatoloqic diseases, N.A. Soter and H.P. Baden, Editors. 1991, McGraw-Hill Inc., Health Professions Division: New York, p. 3-46. Rouabhia, M., Structural and Functional Complexity of the Skin, in Skin Substitute Production by Tissue Enqineerinq: Cl inical and Fundomen ta1 Applications, M. Rouabhia, Editor. 1997, Chapman & Hall: New York, p. 3- 22. Singer, A.J. and R.A.F. Clark, Cutaneous wound healing. The New England Journal of Medicine, 1999. 341 (10): 738-746. Sheridan, R. and R. Tornpkins, Skin substitutes in burns. Burns. 1999. 25: 97-103. Gogia, P. P., Physiology of Wound Heal ing, in Clinicol Wound Mana~emenl, P.P. Gogia, Editor. 1995, SLACK Incorporated: New Jersey, p. 1-12. Baden, H. P., Biology of the Epidermis and Pathophysiology of Psoriasis and Certain lchthyosiforrn Dermatoses, in Pathophvsiology of derrnatoloqic diseases, N.A. Soter and H.P. Baden, Editors. 1991, McGraw-Hill Inc., Health Professions Division: New York, p. 131 -1 58. Sauder, D.N., The Skin as an lmmunologic Organ, in Pathophvsiology of dermatoloqic diseases, N.A. Soter and H.P. Baden, Editors. 1991, McGraw- Hill Inc., Health Professionals Division: New York, p. 101 -1 12. Martin, P., Wound healing - Aiming for perfect skin regenerotion. Science, 1997. 276: 75-80. Kirsner. R.S. and W. H. Eaglestein, The wound healing process. Dermatolosic Clinics, 1993. 1 1 (4): 629-640. Gailit, J. and R.A.F. Clark, Wound repair in the context of extracellular rnatrix. Current Opinion in Ce11 Biology, 1994. 6: 717-725. Mast , B.A., The Skin, in Wound Healinp: Biochemical & Clinicol Aspects. I.K. Cohen, R.F. Dingelmann, and W.J. Lindblad, Editors. 1992, W.B. Saunders Company: Philadelphia, p. 344-355. Cuono, C.B., Physiology of Wound Heaiing, in Cutaneous Wounds, F.J. Dagher, Editor. 1985, Futura Publishing, Co. Inc.: New York, p. 1-22. Hayes, H. , Jr., A review of modern concepts of healing of cutaneous wounds. Journal of Dermatologic Surgery and Oncology, 1977. 3(2): 188- 193.
Pilcher, B. K., et al., Collogenase- 7 and collogen in epidermol repair. Archives of Dermatologie Research, 1998. 29O(Supplement): 537-546. Stenn, K.S. and R. Malhotra, Epithelialization, in Wound Healinq: Biochemical B C[inical Aspects, I.K. Cohen, R.F. Dingelmann, and W. J. Lindblad, Editors. 1992, W.B. Saunders Company: Philadelphia, p. 11 5- 124. Hanson, S.R. and L A . Harker, Blood Coagulation und Blood-Materials Interactions, in Biomateriols Science, B.D. Ratner, et al., Editors. 1996, Academic Press: San Diego, p. 193-199. Tang, L. and J. W. Eaton, lnfIommatory responses tu biomaterials. American Journal of Clinical Pathology , 1995. 103: 466-471. Zitelli, J.A., Synthetic skin. Advances in Dermatology, 1989. 4: 323-342. Wiseman, D.M., et al., Wound Dressings: Design and Use, in Wound Healina: Biochemical & Clinical Aspects, I.K. Cohen, R.F. Dingelmann, and W. J. Lindblad, Editors. 1992, W. B. Saunders Company: Philadelphia, p. 562- 577. Himel, H.N., et ol., Pilot study of a unique film dressing for the treatment of donor site wounds. Journal of Burn Care Et Rehabilitation, 1998. 19: 62- 65. Quinn, K.J., et al., Principles of burn dressings. Biomaterials, 1985. 6: 369- 377. Griswold, J.A., et al., A cornparison of Xeroform and SkinTemp dressings in the healing of skin graft donor sites. Journal of 8urn Care Et Rehabilitation, 1995. 16: 136-140. Horch, R.E. and G.8. Stark, Comprison of the effect of a collogen dressing and a polyurethane dressing on the healing of split thickness skin graft (STSG) donor sites. Scandinavian Journal of Plastic and Reconstructive Hand Surgery, 1998. 32: 407-41 3. Genecov, D.G., et al., A controlled subutmospheric pressure dressing increuses the rate of skin graft donor site reepithelialization. Annals of Plastic Surgery, 1998. 40: 21 9-225. Holcomb, J., LTC, et al., Efficacy of a dry fibrin sealant dressing for hemorrhage control of ter bal 1 istic injury. Archives of Su rgery , 1 998. 1 33 : 32-35. Greenhalgh, D.G., et al. , Multicenter trui l to evaluate the sofety and potential efficacy of pooled human fibrin seolant for the treatment of burn wounds. Journal of Trauma, 1999. 46(3): 433-440. Achauer, B.M., S.R. Miller, and T.E. Lee, The hemostatic effect of fibrin glue on graft donor sites. Journal of Burn Care and Rehabilitation, 1994. 24-28. Lamba, N.M.K., K.A. Woodhouse, and S.L. Cooper, The Chemistry of Polyurethane Copolymers, in Polyurethanes in Biomedical Applications,
A.W. Starkweather and F. Shapiro, Editors. 1997, CRC Press LLC: Boca Raton, p. 5-23. Skarja, G.A. and K.A. Woodhouse, Synthesis and characterization of degradable polyurethane elostomers containing an amino acid-based chain extender. Journal of Biomaterials Science, Polymer Edition, 1998. 9(3): 271 - 295. Lamba, N.M.K., K.A. Woodhouse, and S.L. Cooper, Structure und Physical Characterization of Polyurethanes, in Polvurethanes in Biomedicol Applications, A. W. Starkweather and F. Shapiro, Editors. 1997, CRC Press LLC: Boca Raton, p. 43-89. Painter, P.C. and M.M. Coleman, Structure, in Fundamentols of Polymer Science: An Introductow Text. 1997, Technomic Publishing Co. : Pennsylvania, p. 197-246. Lamba, N.M.K., K. A. Woodhouse, and S.L. Cooper, Bulk Characterization and S tructure-Property Relationships of Polyurethones, in Polvurethanes in Biomedical Applications, A.W. Starkweather and F. Shapiro, Editors. 1997, CRC Press LLC: Boca Raton, p. 43-89. Skarja, G.A. and K.A. Woodhouse, Structure-property relotionships of degradable polyurethane elostomers containing on amino acid-bosed chain extender. Journal of Applied Polymer Science, 2000. 75: 1522-1 534. Lamba, N.M.K., K.A. Woodhouse, and S.L. Cooper, Degrodation of Polyure thanes, in Polvurethanes in Biomedical Applications, A. W. Starkweather and F. Shapiro, Editors. 1997, CRC Press LLC: Boca Raton, p.
Smith, D.C., Adhesives und Sealants, in Biomaterials Science, B.D. Ratner, et al., Editors. 1996, Academic Press: San Diego, p. 319-328. Bense, C.A. and K.A. Woodhouse, Plasmin degradation of fibrin coatings on syn thetic polymer substrotes. Journal of Biomedical Ma terials Research, 1999. 46(3): 305-314. Hantgen, R. R., e t al., Fibrinogen Structure und Physiology, in Hernostasis and Thrombosis - Basic Principles and Clinical Practice, R.W. Colman, et al., Editors. 1987, J.B. Lippincott Company: Philadelphia, p. 269-285. Scardino, M.S., et al., Evaluation of fibrin sealants in cutaneous wound closure. Journal of Biomedical Materiais Research, 1999. 48: 31 5-321 . Wadstrom, J. and A. Tengblad, Fibrin glue reduces the dissolution rate of sodium hyaluronate. Upsala Journal of Medical Science, 1993. 90: 159-1 67. Nakamura, M., M. Hikida, and T. Nakano, Concentration and moleculor weight dependency of robbit corneal epithelial wound heoling on hyaluronan. Current Eye Research, 1992. 11 (10): 981 -986. Hascall, V.C. and T.C. Laurent, Hyaluronan: Structure and physical properties. http: / Iwww.~lycoforum.qr.~p/science/hya~uronan/HAO1 /HA01 E.html, 1997.
Chen, W.Y. J. and G. Abatangelo, Functions of hyaluronan in wound repair. Wound Repair and Regeneration, 1999. 7: 79-89. Manuskiatti, W. and H.I. Maibach, Hyaluronic acid andskin: Wound healing and aging. International Journal of Dermatology, 1996. 35(8): 539-544. Goa, K.L. and P. Benfield, Hyaluronic Acid: A review of its pharmacology and use as a surgical aid in ophthalmology, and its theropeutic potential in joint diseuse and wound healing. Drugs, 1994. 47(3): 536-566. Scott, J.E., Chernical morphology of hyaluronan, in The Chemistry, Bioloqy and Medical Applications of Hyaluronan and its Derivatives, T.C. Laurent, Editor. 1998, Portland Press: London, p. 7-16. Balazs, E.A., The viscoelastic intercellular motrix and control of cell function by hyaluronan, in The Chemistry, Bioloqy and Medicol Applications of Hyaluronan and its Derivatives, T.C. Laurent, Editor. 1998, Portland Press: London, p. 185-204. Heinegard, D., e t al., Hyaluronan-binding matrix proteins, in The Chernistry, Bioloqy and Medical Applications of Hvaluronan and its Derivatives, T.C. Laurent, Editor. 1998, Portland Press: London, p. 1 1 3-1 22. Balazs, E.A. and 2. Darzynkiewicz, The effect of hyaluronic acid on fibroblasts, mononuclear phagocytes and lymphocytes, in 8iolog.y of Fibroblost - 4th Siprid Juselius Foundation Symposium, finlond, E. Kulonen and J.M.K. Haramer, Editors. 1973, Academic Press: London, p. 237-252. Bettinger, D.A., B. Mast, and D. Gore, Hyaluronic Acid lmpedes Reepithelialization of Skin Graft Donor Sites. Journal of Burn Care & Rehabilitation, 1996, 17(4): 302-304. Burd, D.A.R., et al., Hyaluronan and wound healing: a new perspective. British Journal of Plastic Surgery, 1991. 44: 579-584. Goldberg, R.L. and 8.P. Toole, Hyoluronate inhibition of cell proliferation. Arthritis and Rheumatism, 1987. 30(7): 769-778. Hayen, W., e t al., Hyaluronan stimulates tumor cell migration by modulating the fibrin fiber architecture. Journal of Cell Science, 1999. 11 2: 2241 -2251. Laurent, C., S . Hellstrorn, and E. Fellenius, Hyaluronan improves the healing of experimental tymponic membrane perforations. Archives of Otolaryngology Head and Neck Surgery, 1988. 1 14: 1435-1 441. Laurent, T.C., The action of hyaluronan on repoir processes in the middle ear, in The Chernistry, ~ ~ O ~ O Q Y and Medical Applications of Hyaluronan and i ts Derivatives, T.C. Laurent, Editor. 1998, Portland Press: London, p. 283- 290. Murashita, T., et al., Acceleration of granulation tissue ingrowth by hyaluronic acid in artificial skin. British Journal of Plastic Surgery, 1996. 49: 58-63. Suzuki, Y. and T. Yarnaqcht, Effects of hyaluronic acid on macrophage phagocytosis and active oxygen release. Agents Actions, 1993. 38: 32-37.
Doillon, C. J. and F.H. Silver, Col lagen-based wound dressing: Effects of hyaluronic acid and fibronectin on wound healing. Biomaterials, 1986. 7: 3- 8. Huang-Lee, L.L.H., J.H. Wu, and M.E. Nimni, Effect of hyaluronan on collagen fibrillar matrix contraction by fibroblasts. Journal of Biomedical Materials Research, 1994. 28: 123-1 32. Bertheim, U. and S. Hellstrom, The distribution of hyuluronan in human skin and mature, hypertrophie and keloid scars. British Journal of Plastic Surgery, 1994. 47: 483-489. West, D.C., et al., Angiogenesis induced by degradation products of hyoluronic acid. Science, 1985. 228: 1324-1 326. West, D.C. and S. Kumar, Hyaluronon und angiogenesis, in The bioloqy of hvaluronan, CIBA foundation symposium. 1 989, p. 187- 1 99. Svee, K., et al., Acute lung injury fibroblast migration and invasion of a fibrin motrix is mediated by CD44. Journal of Clinical Investigation, 1996. 98(8): 171 3-1727. Weigel, P.H., G.M. Fuller, and R.D. Le Boeuf, A model for the role of hyoluronic acid and fibrin in the early events during the inflammatory response and wound healing. Journal of Theoretical Biology, 1986. 1 19: 21 9- 234.
Weigel, P.H., et al., The specific interaction between fibrin(ogen) and hyuluronan: possible consequences in hoemostasis, inflammation and wound healing, in The bioloqy of hvaluronan. 1 989, p. 248-261. Nehls, V. and R. Herrmann, The configuration of fibrin dots determines capillary morphogenesis and endothelid cell migration. Microvascular Research, 1996. 51: 347-364. Le Boeuf, R. D., e t al., Effects of hyaluronic acid and other glycosaminoglycans on fibrin polymer formation, Biochemistry, 1 987. 26(19): 6052-6057. Le Boeuf, R. D., e t al., Human f ibrinogen specif ical ly binds hyaluronic acid. Journal of Biological Chemistry, 1986. 261 (27): 12586-1 2592. Nahas, L.F. and B.L. Swarts, Use of Semipermeable Polyurethane Membrane for Skin Graf t Dressings. Plastic and Reconstructive Sursery , 1 981 . Skarja, L A . and K.A. Woodhouse, Biodegradoble materials for wound dressings, 1998: United States Patent. Skarja, C.A. and K.A. Woodhouse, In vitro degradation and erosion of degradable, seqmented polyurethanes containing an amino acid-based chain extender. Journal of Biomaterials Science, Polymer Edition, 2000 (submitted). Carey, F.A., Spectroscopy, in Orpanic Chemistry, 2nd Edition. 1992, Mc Graw-Hill, Inc.: New York, p. 499-552. Bovey, F.A. and P.A. Mirau, NMR of Polymers. 1996, San Diego: Academic Press, Inc. 459.
Skarja, L A . , persona1 communication. 1999. van Bogart, J.W.C., P.E. Gibson, and S.L. Cooper, Journal of Polymer Science: Polymer Physics Edition, 1983. 21 : 65. Frontini, P.M., M. Rink, and A. Pavan, Journal of Applied Polymer Science, 1993. 48: 2003. Mun k, P., Introduction to macromolecular science. 1989, New York: Wiley- Interscience. 522. Grodzinsky, A.J., R.D. Kamm, and D.A. Lauffenburger, Quantitative Aspects of Tissue Engineering: Basic Issues in Kinetics, Transport and Mechanics, in Principles of Tissue Enpineerinq, R.P. Lanza, R. Langer, and W.L. Chick, Editors. 1997, R.G. Landes Co.: Austin, TX, p. 193-208. Yannas, I.V. and J.F. Burke, Design of an artificial skin. I. Basic design principles. Journal of Biomedical Materials Research. 1980. l 4 ( l ) : 65-68. Schlag, G. and H. Redl, Fibrin sealant: Efficacy, quatity, and safety, in Fibrin Sealont in Operative Medicine, G. Schlag and H. Redl, Editors. 1994, Springer-Verlag: Berlin, p. 2-1 7. Skarja, G .A., et al., Protein and phte le t interactions wi th thermally denatured fibrinogen and cross-linked f ibrin coated surfaces. Biomaterials, 1998, Rubens, F., e t al., Interactions of thermally denatured fibrinogen on polyethylene wi th plasma proteins and plotelets. Journal of Biomedical Materials Research, 1992. 26: 1651 -1 663. Rubens, F.O., et al., Plotelet accumulation on fibrin-coated polyethylene: Role of platelet activation and factor X I / / . Thrombosis and Haemostasis, 1995. 73(5): 850-856. Bense, C.A., The in vitro stobility of cross-linked fibrin and f ibrinl hyoturonan-cooted polyurethanes in the presence of plasmin, Master's Thesis, Department of Chemical Engineering and Applied Chemistry, University of Toronto, 1998. Fromstein, J. and K.A. Woodhouse, The effect of cross-linking time and factor XI / / concentration on the stubility of a fibrin loyer coated onto wound dressings, Thesis, Chemicat engineering and applied chemistry, University of Toronto, 1998. Hoare, K., et al., Identification of hyaluronan binding proteins usine, a biotinylated hyaluronan probe. Connective Tissue Research, 1993. 30: 1 17- 126.
Appendix One:
NMR Spectra
P W B SEQVEKE R e l a x . delay 1 .000 oac Pulmo 03.0 dogree i Acq. t h 3.000 800
width 8000.0 Ht 64 r o p o t i t i o n a
088HRVI 81, 499.8486422 Küz DATA PROCESSXW
TT m i r a 131072 Total timo 4 min, 16 mec
J o a n u m ~runntein P110600AC 62200 Aug 1 7000
PUT&g SBQUENCip
Rilrn. U m l r y 1.000 roc Pula. 83.0 digrmoi Acq. timi 3.000 mec
Width 8000.0 R t 64 reprtitionm
088ZRM H l , 499.8486456 N€lX
DATA PROCBBSIm FT iizs 131072 Total tinu 4 min, 16 eec
Puma s ~ ~ r r e i c ~ Rmlur. delay 1.000 soc Pulse 83.0 dagrose AC^. t u 3.000 uoc Width 8000.0 Ht 64 rmpetitionm
0881LRV. Fil, 499.8486454 Wffr
DATA PROCEBBfNa FT mixm 131072 Total t h 4 m i n , 16 sec
Joanne rramstdn PBûI000D d2199 Aug 1 2000
PULSE BBQITINCB Rmlmc, dmley 1.000 orna
Pulmo 83.0 degroom Aaq. t h 3.000 Bac Width 0000.0 Er
64 rmpmtitionn OBBmVI: Hl, 499.8486454 WETr DATA PROCEBBINQ PT mite 131072 Total tiw 4 min, 1 6 imc
PULSE BEQ-B R m l m c . dalay 1.000 mec mlis 83.0 dagrose Acq. t h m 3.000 0.C Width 8000.0 B r 64 roprtitioni
OBBBRVX BI, 499 .0486409 lCLTX
DATA PROCESSXbW f i mir. 131072 Total t h e 4 min, 16 rmc
' I I
2.5
Joraam Pranmtain blendlb 62191 Aug 1 2000
PIWB s i q r n a i e r
Rmlax. dmlay 1,000 mec Pu119 83.0 daarmer mq. t h 8 3.000 iac Width 8000.0 Hr 64 rmprtitioni
088RRVE Hl, 499.8406370 wi
DATA PROCtSSINo rn i l a m 131072 ~ot.1 tiaw 4 min, 16 i o c
Solvant t DM80 amp p. 35.0 C / 308.1 R
WITY-500 "ultra500*
P a s a SEQWBBXB
Rmlax. dmlay 1.000 .oc mlsm 0 3 . 0 dmgrmmo Acq. th. 3.000 .ma Width 0000.0 Hz 64 rmpmeitionm
088mV1 Hl, 499.8486405 MHz DATA PROCE88Xm
Linm broadmning 1 . 0 H z PT d a o 131072 Total t h 4 min, 16 m e
P a b i i r SEQURùXB ~mlur. dalay 1.000 imc Pulim 83.0 doaraaa Acq. thno 3.000 ioa Width 8060.0 Hz 64 rmpotitioni
088xRVB Hl, 499.8486405 IMs DATA PROCEBBINO tins broadsning 1.0 Hz
FI' oiro 131072 Total thno 4 min, 16 mec
F a s i c SOQtl.SrcE
R O l u t . d @ h y 1.000 8.C
~ u l i a 83.0 digrmai Acq. t h e 3.000 sec
Wldth 8000,O HZ 64 r o p a t i t i o n o
OBBBR- 81, 499.8486405 BUà+
DATA PROCBBBIWQ Lino broadening 1 . 0 Hz
FT m i r e 131072 T o t a l t h e 4 min, 16 arc
P a s r ~ W K N C E usla%. delay 1.000 mec mlim 8 3 . 0 dagrseo Acq. t h 3.000 s e c wid th 8000.0 Ba 64 r m m t i t i o n n
OBSBRVZ El, 499.0486418 iUSS
DATA PROCESGIm
Linm broadrning 1 . 0 Hz PT iirm 131071
~ o t a l t h 4 min, 16 sac
P m 8 s.qwmcic Rmlmc. daliy 1.000 mao Pulm 83.0 Urgrems ACq. t h 3.000 €WC
Width 8000.0 t f t
64 rrpmtitionm o~smlrm Hl, 499.8486502 MHz DATA PROCISBIW Lin. broadmning 1.0 Hz fi m i . . 131072 ~otml tiar 4 min, 16 mec
Appendix Two:
DSC Spectra
Sarnp le: PCL2000-A Sire: 8.4000 mg Method: 30ANNA DSC
file: J O A 7 D S C Ope~ator: HFG Run Date: 31-Aug-O0 09: 24
Temperature (OC) General V4.1C DuPont 2100
DSC
PCL2000-B 2ND RUN QUENCHED
Temperature (OC) General V 4 . X DuPont 2100
Santple: 6008 Size: 28.0000 mg Method: JOANNA DSC
File: 30.22 D s C O P B P ~ ~ O P : HFO Aun Date: 31-Aug-O0 16: 14
Temperature (OC) General V 4 . 1 C DuPont 2100
- a . a P O C O € N U E 1 0 d PI O m c n r o
Simple: 2-A Sire: 8.0000 mQ Msthod: JOANNA DSC Comment: -140°C TO 250%. 15C/MiN. N2
DSC File: J0.04 Operator: HFG Run Date: 29-Aug-O0 30: 25
Temperature (OC) General V 4 . X DuPont 2100
Sample: 3-A Size: 10 .O000 mg Methad: JOANNA DSC Comment: -140°C TO 250%. 15C/MIN. N2
DSC File: &.O8 Operator: HFG Aun Date: 30-Aug-00 OB: 40
Temperature (OC) General V 4 A C DuPont 2i00
Appendix Three: Development of Three-Dimensional Foams Using
Biodegrada ble Polyurethane Blends - - .. .
Introduction: For many tissue engineering applications, the use of biodegradable three-
dimensional porous constructs i s highly desirable. These materials could provide
the mechanical and structural support necessary for preliminary ce11 and nutrient
infiltration. As tissue formation proceeds, the scaffold i s no longer needed for as
much support, and can therefore begin to degrade, transferring i t s Load to the
newly developed tissue. In order for this approach to work, the material must be
highly porous, possessing a high degree of interconnectivity between pores, as well
as a high surface to volume ratio, and pore sizes large enough to allow for cells to
pass easily through[A3-2, A3-31. For applications where the cells are not being
seeded prior to implantation, pore sizes ranging from 20-2OO~im have been found
to be optimal for soft tissues[A3-1, A3-2, A3-41.
In this appendix, some very prelirninary work i s presented, whereby three-
dimensional scaffolds were created out of all four blends. The foams were cast
using a "solvent-casting, particulate leaching" method, which has been used by
several investigators previously to form porous materials[A3-5, A3-6, A3-71. These
preliminary investigations were performed simply to demonstrate the ability to
fabricate three-dimensional structures using the btends developed in this thesis.
Materials and Methods: The same materials used in fabricating the blends (chapter three) were used for
the three-dimensional foams. Reagent grade sodium chloride (Aldrich, Milwaukee,
WI) was used as the particulate matter. Prior to use, the NaCl was ground using a
porcelain bal1 mill, and then passed through a sieve in order to ensure that all salt
particles were no iarger than 150pm in size. Using this qinding process, over 90%
of the salt was recovered.
In designing the method for making foams, it was decided that small, square
polytetrafluoroethylene (PTFE, Teflonm') casting, dishes should be used. Dishes
that are 8 cm x 8 cm x 1 cm were designed. This size was chosen as the foams
made through this process would need to be big enough to allow for mechanical
tensile tests to be performed eventually, yet small enoush that excessively large
amounts of polymer need not be used. The new dishes were fabricated to order,
ensuring that the bottom of the dishes was very flat. This was an important
factor, as the commercially available Teflon dishes tend to be round and have
concave bottoms. I t was a concern that surface tension effects caused by this
concaveness may have affected the formation of a uniform foam across the entire
length.
To fabricate the three-dimensional structures, standard solvent casting-particdate
leaching methods were adapted from those of Mikos et al.[A3-61. Widmer et
al.[A3-71 found that increasing the percent salt (by weight) in the casting solution
resulted in increased porosities and interconnectivities. Pore size can also be
affected, by controlling the salt particle size. Widmer e t al. found that salt mass
percentages of at least 70% were required to form foams that were uniform and
did not possess a "skin" of polymer on top. Based on these results, salt
percentages of 85% and 95% were selected for this study. The most significant
adaptation made was a decision that total mass of polymer and salt should be held
constant, rather than just the mass of polymer. If the polymer mass had be held
constant, the size of the foams would V a r y significantly more due to the wide
changes in salt masses needed for the hisher salt percentages.
In order to determine the total mass that should be used, the dishes were filled
with 30 mL of water, and then pure salt was added unti l the dish was full. The
mass of salt required to fil1 the dish was 33 g. Therefore, this mass was rounded
off and set as the mass of salt to be used for 95% salt foams. The total mass used
was thus chosen as 35 2. Table 3-1, below indicates the masses of salt and of each
polyurethane used to make the various foams:
Foam % Salt Salt Mass Mass Mass No. (by mass) Mass (g) PhelPCL2000 (g) PhelPE0600 (8) PhelPEObOO (g)
Table 3-1: Compositions of the six preliminary foams created using solvent casting-particutate Leaching methods.
The necessary masses of polyurethane were weighed out, and dissolved in 25 mL
chloroform. Following dissolution, the chloroform solutions were filtered to
remove any impurities. In the cases of foams 1 and 4 through 6. the
predetermined amount of salt was laid flat, and evenly distributed in the casting
dishes. The polyurethane-solvent solution was then poured over the salt mixture.
For foarns 2 and 3, in order to determine the effect on casting, the salt was
poured into the solution, and the solution well mixed prior to casting. Once al1
the casting dishes had been filled, they were covered to prevent impurities from
f a h g in, and le f t to evaporate for 48 hours. The foams were then transferred to
a vacuum oven, and dried a t roorn temperature for 48 hours, t o ensure that al1 the
chloroforrn had been removed from the scaffolds. The foams were then soaked in
water under light agitation, in order to remove the salt from the scaffolds. This
rinsing process was carried out for a total of 4 days, and the water was changed
three times a day.
Re& and Discussion: The six materials formed were visibly porous. Foam #2 showed a thick skin on top,
a property often seen with materials cast in this way. Foam #3 had a very rough
surface, possibly caused by air bubbles present during casting. The other four
materials al1 appeared to be fairly uniform. From these results, it was determined
that laying the salt down in the casting dish, and then pouring the polyurethane
solution over the salt is a better, more consistent casting method. This agrees
with observations made during casting that if the salt is mixed into the solvent
prior to casting, a lot of salt sets Left behind, stuck to the sides and bottom of the
flask. It i s possible that some polyurethane may also have been retained with the
salt. Thus, the salt to polymer mass ratios could have, and likely did, change.
Based on these results, salt should be pre-laid in the casting dishes for any future
experimen ts.
After leaching, and drying, the materials s t i l l felt grainy, indicating that not al1 of
the salt has been removed from the scaffolds. However, the Longer the foams are
soaked in water in an attempt to leach the salt, more of the polyurethane blends,
particularly the PEO-containin9 components, wi l l be degraded and removed. The
results presented earlier in the thesis indicate that the blends degrade rapidly
early on, and then level off, degrading much slower after approximately two
weeks time. If this degadation profile i s desired and i s considered in the design
of the three-dimensional scaffolds, Longer leaching periods could affect the
results. The foams could be partially degraded prior to implantation. Also, since
it was hypothesized that the PEO-containing segments are degraded first, the
composition of the blends could be significantly different from where they started.
As with the blends thernselves, since the polyurethanes were filtered prior to
casting, the actual ratios of PCL-polyurethane to PEO-containing material may
differ from the expected values. Thus, analytical techniques such as gel
permeation chromatography (GPC) must be used to determine the final
compositions of the foams. In addition, the foams wil l need to be characterized,
in order to determine the pore size, porosity and interconnectivity of the pores.
This can be determined using mercury intrusion porosimetry and scannine, electron
microscopy .
The blends created using these methods feel quite strong to the touch. They can
be stretched a litt le, but do not deform permanently when pulled by hand. I t i s
difficult to compare these materials to the flat blend films based on these
observations, since the foams are significantly thicker (approximately 3-4 mm in
thickness). I t is expected that the foams are stronger, as the walls of the pores
interconnect different points in the materiai, acting similarly to cross-links.
However, details about the difference in strength, and the effect of composition
or salt concentration on these values, wil l need to be evaluated through
comparative mechanical tensile testing.
Conc/usions and Recommendatim: These results demonstrated that three-dimensional structures can be formed using
the biodegradable polyurethane blends developed in this work. However, in order
for these structures to be used, the casting methods will need to be perfected. A
salt concentration should be chosen, and once reproducible blends are produced,
they will need to be characterized not only for their three-dimensionality, but also
for their chernical composition. Degradation studies and mechanical tensile
testins wi l l also be necessary.
The inherent porosity of the blended materials, as seen in the degradation study
(chapter three) , may be advantageous in the formation of three-dimensional
foams. The small pores formed by the material itself may provide interconnecting
channels between the pores created by the salt crystals. Upon completion of
these studies, a viable three-dimensional biodegradable scaffold could be
developed, which may be used in a wide variety of soft tissue engineering
applications.
Re ferences
A3-1. Yannas, I.V. and J.F. Burke, Design of an artificiel skin. I . Basic design principles. Journal of Biomedical Materials Research, 1980. l 4 ( l ) : 65-68.
A3-2. Kadiyala, S. , H. Lo, and K. W. Leong, Formation of highly porous po[ymeric fooms wi th controlied release capability - A phase-separation technique, in Tissue Enqineerinq Methods and Protocois, J.R. Morgan and M.L. Yarmush, Editors. 1999, Humana Press Inc.: New Jersey, p. 57-65.
A3-3. Garrido, L., Nondestructive evaluation of biodegradable porous matrices for tissue engineering, in Tissue Enqineerinq Methods and Protocois, J . R. Morgan and M.L. Yarmush, Editors. 1999, Humana Press Inc.: New Jersey, p. 35-45.
A3-4. Whang, K., C.H. Thomas, and K.€. Healy, A novel method to fabricate bioabsorbable scaffolds. Polymer, 1995. 36(4): 837-842.
A3-5. Ma, P.S. and R. Langer, Fabrication of biodegradable polymer foams for cell tronsplontotion and tissue engineering, i n Tissue Enqineerinq Methods and Protocois, J .R. Morgan and M.L. Yarmush, Editors. 1999, Humana Press Inc.: New Jersey, p. 47-56.
A3-6. Mikos, A. G. , et al., Preparation and chorocterizafion of poly(L4actic acid) foams. Polymer, 1994. 35(5): 1068-1077.
A3-7. Widmer, M. S., et al., Manufacture of porous biodegradable polymer conduits by an extrusion process for guided tissue regenerotion. Biomaterials, 1998. 19: 1945-1955.
Appendix Four:
Preliminary I n Wvo Evaluation of the Fibrin Coatings
introduction In designing a fibrin-coated polyurethane wound dressing, the fibrin incorporation
was expected to improve adherence to the wound site and promote faster healing.
Therefore, once the coatings had been developed, a preliminary in vivo
investigation was performed in order to determine whether the presence of the
coating affects the healing process. In addition, the biodegradable polyurethane
blends and parent polymers had never before been tested in vivo. Therefore,
some preliminary observations of the effects of these materials being present
during wound healing were made. The results are presented in this appendix.
Materials and Methods
Materials Four materials were tested using this in vivo model: the pure Phe/ PCL2000, blend
1 , fibrin-coated blend 1, and ~ p s i t e v s m i t h + Nephew). Blend 1 was selected
based on i t s ease of suturability, as discussed earlier in chapter three. In addition,
in vivo studies are expensive, and therefore, one test material and one parent
material were seen as an acceptable representation for a preliminary study of this
nature.
Methods For these experirnents, a new five wound fut[-thickness healing model was utilised,
developed in collaboration with the Defense and Civil lnstitute of Environmental
Medicine (DCIEM). Five 0.5 kg Sprague-Dawley rats were used in these
experiments. Two experiments were performed, the first, a 48-hour study
involving two rats, used all four materials listed above. Prior to administration of
the dressings, they were sterilized by dipping them in 100% isopropyi alcohol, and
then rinsing them in sterile distilled water. The second experiment was carried
out to 28 days, using three rats, and al1 materials except opsite'. These materials
were prepared usine, ultraviolet (U.V.) sterilization. The in vivo protocols were
the same for both experiments, as described below.
The layout of the wounds on the rats' backs i s illustrated in Figure 4-1. Prior to
operating, the rats were anaesthetized using 2% Halothane gas. Injections of
0.025 mg of Terngesic were also administered in order to prevent the animals from
feeling pain during and after the experiments.
Figure 4-1: Five wound modet used for testing polyurethane dressings. Each wound was 1 cm x 1 cm in size.
Throughout the procedures, the rats were kept under anaesthesia using a mask.
Once anaesthetized, the rats' backs were shaved and anitbacterialized. A
template of the five wounds was used to "spray paint" the pattern ont0 the backs
of the rats. Next, full thickness incisions were cut along the edges of the squares
using a scalpel, forceps and scissors. The square pieces of skin were removed, and
any excess blood was absorbed using gauze. Next, square pieces of the various
dressings the same size as the wounds were sutured into place. The dressings
were tucked under the skin at the wound edges in order to ensure good contact
between the wound beds and dressings (Figure 4-2). Anywhere from four to seven
sutures were utilised, in order to ensure that the dressings were held firmly, and
the contours of the wound edges rernained square. Controi wounds were left
uncovered, and were not sutured.
Following the surgery, the rats were placed in an oxygen recovery chamber until
recovery. Temgesic was re-administered twice daity for the first 2 days of the
experiments. Foliowine, 48 hours for the first study, and 28 days for the second.
the rats were euthanized using injections of T-61 euthanising solution.
After the first study, the wounds were grossly analysed, and measured for wound
bed area. For the second study, the rats' wound beds were surgically excised
(following the same protocols as for the wound creation, onty cutting out an area
slightly Larger than the originaL wound size), fixed in 10% formalyn soiution, and
sent for histologie processing using haematoxlyn phloxine saffron (HPS) stain.
Sutures
/ Wound drcssing
Figure 4-2: Cross-sectional view of the wound dressing placement within the wound bed.
Resuits and Discussion
Preliminarv Ex~eriment The first in vivo study, conducted for a 48-hour period, was used in order to
ensure that this new five-wound model was effective, and to work out any
difficulties associated with the model. Durinp surgery, it was noticed that
dependin5 on where each wound was situated on the rat's backs, there were
differences in skin thickness and amount of bleeding. The differences in skin
thickness played a role in how easily suturable the dressings were. In addition,
differences in bleeding as well as position could lead to different rates or qualities
of healins. Thus, it was determined that in future, the positions of the dressings
should be different on different anirnals so that the location effects can be
determined or minimized i f need be.
Also, since the wound dressings are supposed to be semi-permeable, we did not
want to wrap the rats' backs with gauze or other protective materials that might
interfere with or affect the healing process. Following surpery, the rats were
returned to their cages, however, which were filled with wood chips, food, and
other substances that could contaminate the wounds i f the rats rolled ont0 their
backs in their cages. Thus, a dome-like construct (F iy re 4-3) was developed to
protect the wounds from these external factors. The construct consisted primarily
of a thin sheet of plastic approximately 10 cm Lon2 and 6 cm wide. The plastic
was thin enough to be rolled, while s t i l l being rigid enough to retain the roiled
shape. ~ e f i x ' dressings (SCA Molnlycke Ltd.), a water vapour permeable polyester
coated with acrylic adhesive on one side, were used to cover the two ends of the
dornes. Small holes were punched in the four corners of the rigid plastic forming
the dome, and ~ o b a n self-adhesive tape (3M) was passed through Ehese holes.
Prior to placing the domes on the rats, the units were autoclaved to sterilize
them. The ~ o b a n 9 a p e was then wrapped around the rats' shoulders and Legs to
secure the dome in place. This dome was quite efficient at keepinsj foreign
objects away from the wounds. Some problems were seen, however, with ulcers
forming under the rats shoulders where the ~oban ' dug into the armpits. Also,
condensation appeared to be occurring within the dome, requiring tiny air holes to
be punctured into the dome's plastic.
CobanB straps
Figure 4-3: Oome construct designed to protect dorsal wounds from foreign debris on rats. The CobanS self-adhesive straps are wrapped around the arms and legs of the rat in order to stabilize the dome.
Following 48 hours, the wound sizes were measured to determine the effect. if
any, of dressings on wound bed area. These results are summarized in Table 4-1.
below. The appearances of the wounds were also quali tatively compared.
Rat No. Wound Number Wound Dressing Final Wound Bed Area (cm x cm)
Fibrin coated Blend 1
1 5 None (cantrol) 1.5 x 1.1
Blend 1 0.85 X 1 .O
Rat No. ' Wound Number Wound Dressinq Bed 1 I
Area (cm x cm)
opsi te"
1
2 1 3
2 5 None (control) 1.3 x 1.1
1
Fibrin coated i
Table 4-1: Results from the first in vivo study, indicating effect of dressing type (or lack thereof) on the final wound area of a full-thickness wound following 48 days. Wound nurnber corresponds to the numbers assigned in Fisure 4-1.
From these results, i t can be seen that the wound areas tended to decrease in size
in al1 cases but the controls, where the areas actually increased. This i s likely due
primarily to the fact that the wounds containing dressings were sutured, thereby
forcing the wound edges to rernain close together. The control wounds, however,
were lef t unsutured and therefore had no support holding them closed.
I Biend 1 0.9 x 1.0 I
Qualitative assessrnent of the wounds' appearances demonstrated varying
decrease of vascularization and granulation tissue formation. The wounds treated
with blend 1, both uncoated and fibrin coated, showed more vascularization than
the opsite" and Phe/PCL2000 treated wounds. The control wounds also
demonstrated numerous blood vessek. The wounds under Opsite' and
PhelPCUOOO also appeared to be significantly more moist than the control wounds
and those treated with fibrin coated blend 1.
Second Investiaation
The second study, extended over 28 days, provided some insight into the effects of
these various dressings on cutaneous wound healing in the rat. By the end of the
28 days, six of the fifteen wounds had healed completely. The wound dressings
fell off nine of the wounds pnor to 28 days. For al[ three rats, the final
appearance of the wounds was assessed, and the wounds ranked in order from best
healed (1) to worst healed (5). These ranking were made blindly, without
knowing which wound had had which dressing (although the controls were
distinguishable as they contained no sutures). The results of these assessments
are summarized in Tables 4-2 through 4-4.
Wound Number Dressinq Type Rank
1 None (control) 2
2 Blend 1 4
3 Fi brin -coated Blend 1 1
5 None (control) 2 - - - - - -- -- -- - --
Table 4-2: Qualitative ranking of the wounds on rat $3, rated from best healed to worst healed.
Wound Number Dressing Type Rank - - - -
1 P he / PCL2000 5 --
2 None (control) 2
3 Blend 1 2
I
5 Blend 1 1 --
Table 4-3: Qualitative ranking of the wounds on rat $4, rated from best healed to worst healed.
Wound Nurnber Dressing Type Rank
1 Fibrin-coated Blend 1 3
3 None (control) 1
4 Blend 1 5 t
5 , Fibrin-coated Blend 1 , 2
Table 4-4: Qualitative ranking of the wounds on rat #5, rated from best healed to worst healed.
From these results, it was encouraging to see that the fibrin coated blend 1
dressings were always ranked either first or second for gross quality of healing.
The uncoated blend 1 ranked well on rat #4, but poorly on the other two rats.
Ranking was very difficult for rat #4, as al1 wounds were very similar in
appearance except for wound number 5, which was by far the worse. This rat
displayed significantly worse healing than the other two animals, likely due to
complications during recovery following surgery. Its wounds contained large
amounts of moist, pinkish-red granulation tissue following 28 days. Most of the
well-healed wounds on the other rats, however, were dry, and exhibited white
tissue where re-epithelialization had taken place.
More dressings rernained in place on rat ff4 at the end of the study as well.
Because of how the dressings were originally anchored beneath the skin
surrounding the wound, re-epithelialization may not have been possible until the
dressings had fallen off. Alternatively, if the keratinocytes migrated under the
wound dressings that did fall off early, the process of re-epithelialization could
have caused the removal of the dressings. Either way, for some reason rat # 4 did
not appear to show the same re-epithelialization that the other rats did. This
again may be due to the complications experienced by the rat during surgery.
The other rats had Lost most of their dressings within the first 15 days of the study.
By 28 days, many wounds appeared to be completely healed. Assessment at an
earlier time point, or use of a mode1 that possesses slower healing rates would be
appropriate for future investigations.
Histology Resul ts :
In addition to gros analysis of wound appearance, the wound beds were sent for
histology. A low rnagnification (4x) histolopy image of rat skin is shown in Figure
4-4, below, in order to demonstrate locations of the various parts of the skin. This
image i s of a control wound following 28 days of healingl, as no samples of un-
injured tissue were taken for histology.
Haematoxlyn phloxine saffron (HPS) stain allows for easy differentiation between
collagen and c e k . The tightness of collagen packing, and density of ce11
populations were therefore analysed for the various wounds, and cornparisons
were made. Figure 4-5, below demonstrates the papillary dermis in rat # 3 for i t s
wounds numbered 1 through 4. From these images, it i s evident that the collagen
i s tightest in image A, representing a control wound, and i s much looser for images
C and D, representing the fibrin coated blend and Phe/PCL2000. Therefore, even
though throuph gros appearance the fibrin-coated blend 1 covered wound
appeared to have healed best, i t s collagen deposition is not as tight and neatly
packed as for the control and uncoated blend 1 wounds. The collagen layout i s
similar to that of the PhelPCL2000- covered wound, a material that consistently
demonstrated poorer healing through gross ranking.
Figure 4-4: HPS stained cross-section of rat skin at 4x masnification. The pink and purple near the top of the image represent the predorninantly cellular epidermis. HPS stains cells purptelpink, and collagen an orange-yellow. The purple dots within the collagen represent fibroblasts and other cellular components. The collagen closest to the epidermis i s more closely packed, and represents the papillary dermis, while the reticular dermis i s less wetl packed. The Sap between the papillary and reticular dermises i s simply an artefact created during the histoiogy sarnple preparation.
Al1 four images in Figure 4-5 demonstrate cells present within the collagen bundles
in the dermis. The number of cells present in the control wound appears to be
greater than in the other three sites. In all four images, blood vessels can be
seen. These features appear as dark pink oval-shaped clusters. Evidence of
vascularization i s important in confirming that the wound heaied normally, with a
supply of cells and nutrients to the wound bed. Thus, none of the dressings
prevented vascularization of the wound site. In addition, al1 four images
demonstrate re-epithelialization of the wounds has occurred. The thickness cf the
epidermal layer varies, but since different regions of skin on the rat may Vary in
thickness, this parameter should not be directly compared.
4 - * / - - . - * C : .- # - . a - . - - ,'f '. . , . , . - / -1 .- - -. - -- - ..- - - , .- . --- * l y - - . . ,... -
A.: -.- -- .x'-- -- -- !:A - - y q17 . . - . i-., -- - .-- d - -.- - . : , , ..- .b --. <
- 1 - _ - . Z -
Figure 4-5: Representative HPS stained histology samples of the papillary dermis and epidermis portions of full-thickness wounds heaied in the presence of various wound dressing for 28 days. A) Control (open wound), 0) Blend 1, C) fibrin coated Blend 1, 0 ) Phe/PCL2000. Images were taken at 10x magnification.
Figure 4-6, below, shows the deep reticular dermis sections of the same four
wounds shown above. In these images, blood vessels are seen in only the control
wound and the fibrin coated blend 1 treated wound. This indicates that
vascularization within the deeper regions of the dermis may have been prevented
throuph the use of uncoated blend 1 and PhelPCL2000. It may be that the
addition of a fibrin coating to the blend dressing helped to irnprove the healine,
response, as was hypothesized.
In observing the images in Figure 4-6, it i s evident that the control wound
demonstrated the most cellular infiltration, blood vesse1 formation, and collagen
deposition. The fibrin-coated blend 1 showed the next best healing, as assessed
through these factors. This improved deep vascularity, cell presence and collagen
deposition may help to explain why this wound was qualitatively assessed as the
best-healed of the five wounds. This reasoning would agree with the Lack of much
visible collagen deposition in the Phe/PCL2000 covered wound, which was the
worst healed wound as assessed quali tatively .
Figure 4-6: HPS stained histology images of the deepest regions of the dermis. Samples taken from full-thickness wounds following 28 days of healing that were treated with: A) no dressing, B) biend 1 , C) fibrin coated btend 1, and D) PheIPCU000. Images taken at 10x rnagnification.
Based on these results, it appears that the presence of a fibrin coating on a
degradable polyurethane blend can be used to alter, and therefore possibly
improve, the healing response. In addition, the blend 1 dressings appeared to
perform better than pure PhelPCLZOOO overall. Thus, the use of a blend, and
manipulation of blend composition, could also be used as a means of affecting
wound healing.
Cmc/usims and Recommendat.ons:
Both the fibrin coating and the biodegradable polyurethanes were tested for the
first time in vivo durine, this investigation. A five-wound fut\-thickness wound
model on rats and a dome for protection of the dorsal wounds were developed.
From these very preliminary results, the fibrin-coated blend 1 appeared to
perform better than either the uncoated blend and its parent polyurethane
Phe/PCL2000. These results are based both on gros qualitative assessment, and
histological appearance of collagen layout and ce11 composition.
In future, an animal model which heals more slowly and has thicker skin than rats
should be utilised. The heahg of pig skin is similar to that of human skin, and
would therefore provide a more accurate representation. A model that wil l mirnic
the proposed function of the biomaterials (Le. skin graft donor sites) should be
used to increase the accuracy of the results. Futher in vivo investigations should
also focus not only on the fibrin coating, but also on the fibrin/HA coating. In
addition, the 50:50 blends should also be tested, as they possess a higher degree
of porosity .