modification of biodegradable polyurethanes by
TRANSCRIPT
VYSOKÉ UČENÍ TECHNICKÉ V BRNĚBRNO UNIVERSITY OF TECHNOLOGY
FAKULTA CHEMICKÁÚSTAV CHEMIE MATERIÁLŮ
FACULTY OF CHEMISTRYINSTITUTE OF MATERIALS SCIENCE
MODIFICATION OF BIODEGRADABLEPOLYURETHANES BY BIOLOGICALLY ACTIVESUBSTANCES
MODIFIKACE BIODEGRADABILNÍCH POLYURETHANŮ BIOLOGICKY AKTIVNÍMI LÁTKAMI
DIZERTAČNÍ PRÁCEDOCTORAL THESIS
AUTOR PRÁCE Ing. VOJTĚCH KUPKAAUTHOR
VEDOUCÍ PRÁCE doc. Ing. LUCY VOJTOVÁ, Ph.D.SUPERVISOR
BRNO 2015
Vysoké učení technické v BrněFakulta chemická
Purkyňova 464/118, 61200 Brno 12
Zadání dizertační práce
Číslo dizertační práce: FCH-DIZ0107/2014 Akademický rok: 2014/2015Ústav: Ústav chemie materiálůStudent(ka): Ing. Vojtěch KupkaStudijní program: Makromolekulární chemie (P1422) Studijní obor: Chemie makromolekulárních materiálů (1405V003) Vedoucí práce doc. Ing. Lucy Vojtová, Ph.D.Konzultanti:
Název dizertační práce:Modifikace biodegradabilních polyurethanů biologicky aktivními látkami
Zadání dizertační práce:1) Vypracování literární rešerše na téma biodegradabilních polyurethanů a jejich modifikací biologickyaktivními látkami2) Příprava amfifilních biodegradabilních polyurethanových elastomerních filmů s různými poměry hydrofilnía hydrofobní složky a různým izokyanátovým poměrem bez použití rozpouštědla3) Chemicko-fyzikální charakterizace připravených polyurethanových filmů4) Modifikace a charakterizace vybraných PU filmů biologicky aktivními nanočásticemi s povrchovouúpravou zajišťující dobrou distribuci v PU matrici 5) Shrnutí výsledků a závěr
Termín odevzdání dizertační práce: 20.7.2015Dizertační práce se odevzdává v děkanem stanoveném počtu exemplářů na sekretariát ústavu a velektronické formě vedoucímu dizertační práce. Toto zadání je přílohou dizertační práce.
- - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - -Ing. Vojtěch Kupka doc. Ing. Lucy Vojtová, Ph.D. prof. RNDr. Josef Jančář, CSc.
Student(ka) Vedoucí práce Ředitel ústavu
- - - - - - - - - - - - - - - - - - - - - - -V Brně, dne 20.9.2014 prof. Ing. Jaromír Havlica, DrSc.
Děkan fakulty
3
Abstract:
Presented dissertation thesis is focused on novel preparation of biodegradable
polyurethanes (PUs) and their modification by biologically active cellulose nanocrystals.
Literary review deals with current state of bioresorbable PUs used in tissue engineering.
Examples of prepared PU elastomers, scaffolds and injectable PUs, together with
biodegradation pathways to non-toxic products are summarized. The last part of the literary
review is targeting on nanocellulose, which has gained much attention for the use as
biomedical material due to its remarkable physical (high specific surface area, mechanical
reinforcement) and biological (biocompatibility, biodegradability and low toxicity) properties.
Experimental part presents characterization of biodegradable amphiphilic polyurethane
films (bio-PUs) synthesized by solvent free polyaddition reaction of hydrophilic
poly(ethylene glycol) (PEG) and hydrophobic poly(-caprolactone) (PCL) as macrodiols with
hexamethylene diisocyanate. Prepared bio-PUs were characterized on one hand by means of
different PEG/PCL ratio and on the other hand by changing the isocyanate ratio between
NCO/OH groups. Abrupt enhancement of mechanical properties was observed when
PEG/PCL weight ratio was equal to or less than 20/80 and was ascribed to the PCL ability to
form crystalline domains. The increasing amount of PEG promoted the ability of bio-PUs to
absorb water and enhance the rate of hydrolytic degradation. Whereas, reducing the ability of
bio-PUs to absorb water and prolonged time of hydrolytic degradation was achieved with
increasing the crosslink density by enhancing the isocyanate ratio. The last part deals with
novel solvent free preparation of nanocomposite utilizing bio-PU as a matrix and cellulose
nanocrystals either neat or surface grafted by PEG. Structural analysis demonstrated that the
presence of rod-like nanoparticles causes the immobilization of the PU chains in matrix
resulting in increased stiffness and rigidity of bio-PU/cellulose nanocomposite.
By adjusting the PEG/PCL ratio, the amount of isocyanate or the presence of nanofiller,
the novel bio-PU material with desirable mechanical (toughness, flexibility) and physical
(swelling, degradation) properties can be obtained. Prepared solvent free bio-PUs may
advantageously be used in regenerative medicine for soft tissue regeneration (e.g. as vascular
grafts).
Keywords:
Polyurethane, poly(ethylene glycol), poly(-caprolactone), nanocellulose, nanocomposite,
tissue engineering
4
Abstrakt:
Předkládaná dizertační práce se zabývá novým způsobem přípravy biodegradabilních
polyuretanů (PU) a jejich modifikací biologicky aktivními celulózovými nanokrystaly.
Literární rešerše se zaměřuje na bioresorbovatelné PU v tkáňovém inženýrství. Shrnuje
příklady těchto PU elastomerů, skafoldů (nosičů buněk) i injektovatelných PU společně se
způsoby biodegradace na netoxické produkty. Poslední část je zaměřena na nanocelulózu,
která si získala pozornost díky svým pozoruhodným fyzikálním (velký specifický povrch,
mechanické vlastnosti) a biologickým (biokompatibilita, biodegradabilita a nízký toxicita)
vlastnostem jako materiál pro biomedicínu.
V experimentální části byly charakterizovány amfifilní biodegradovatelné polyuretanové
filmy (bio-PU) syntetizované bez použití rozpouštědla polyadiční reakcí z hydrofilního
poly(ethylenglykolu) (PEG) a hydrofobního poly(-kaprolaktonu) (PCL) jako makrodiolů
společně s hexamethylen diizokyanátem. Připravené bio-PU filmy byly charakterizovány pro
různé poměry jak mezi PEG/PCL, tak i mezi NCO/OH reagujícími skupinami (izokyanátový
poměr). Bio-PU filmy projevily markantní nárůst mechanických vlastností při hmotnostním
poměru PEG/PCL rovnému nebo menšímu než 20/80 díky vzniku krystalických domén PCL.
Přítomnost PEGu zvyšovala schopnost bio-PU filmu absorbovat vodu i urychlila jeho
hydrolytickou degradaci. Oproti tomu nižší absorpční schopnost a delší čas hydrolytické
degradace materiálu způsobil vyšší izokyanátový poměr, a tedy i vyšší síťová hustota. Třetí
část práce se zabývá přípravou polyuretanových nanokompozitů unikátní metodou bez použití
rozpouštědla za využití bio-PU matrice a celulózových nanokrystalů buď nemodifikovaných,
nebo povrchově roubovaných PEGem. Strukturní analýza prokázala, že přítomnost
tyčinkovitých nanočástic způsobuje imobilizaci polymerních segmentů, v důsledku čehož se
zvýšila tuhost a křehkost materiálu.
Nastavením vhodného poměru mezi PEG/PCL, množstvím izokyanátu, či přídavkem
modifikovaného nanoplniva může být bio-PU materiál „ušit na míru“ s vhodnými
mechanickými (houževnatost, tažnost) a fyzikálními (botnání, degradace) vlastnostmi. Díky
přípravě bez použití rozpouštědla by mohly být připravené materiály využity v regenerativní
medicíně jako např. cévní štěpy.
Klíčová slova:
polyuretan, polyethylenglykol, poly(-kaprolakton), nanocelulóza, nanokompozit, tkáňové
inženýrství
5
KUPKA, V. Modification of Polyurethanes by Biologically Active Substances. Brno: Brno
University of Technology, Faculty of Chemistry, 2015. 98 p.
Supervisor doc. Ing. Lucy Vojtová, Ph.D.
Declaration
I declare that my Doctoral Thesis was worked out by myself and that used references
are quoted correctly and fully. The content of the above mentioned thesis is the property of
BUT Faculty of Chemistry and can be used for commercial purposes only with the
supervisor’s and dean’s consents.
……………………….
autograph of author
Acknowledgement:
I would like to thank my supervisor doc. Ing. Lucy Vojtová, Ph.D. for her professional
guidance and prof. RNDr. Josef Jančář, CSc. for providing me the working conditions. I
would also like to thank Assoc. Prof. Qi Zhou for supplying cellulose nanocrystals, Mohd
Farhan Ansari for DMA measurements, doc. RNDr. Miroslav Šlouf, Ph.D. for acquiring TEM
images, Mgr. Zdenka Fohlerová, Ph.D. for cell response analysis and Ing. Radka Bálková,
Ph.D. for discussions regarding structural analysis.
I gratefully appreciate the help of all my colleagues, present and past, at Faculty of
Chemistry, BUT, at the Department of Materials Science for making my stay at University a
pleasant one.
Finally, yet importantly I would also like to thank my parents for their lifelong support and
my dear wife Hanča.
6
List of Contents:
1. INTRODUCTION .......................................................................................................... 8
2. THEORETICAL PART ............................................................................................... 10
2.1. Chemical structure of polyurethanes ...................................................................... 10
2.2. Bioresorbable polyurethanes .................................................................................. 12
2.3. Feedstocks for synthesis of bioresorbable polyurethanes ...................................... 13
Hard segments .................................................................................................. 14 2.3.1.
Soft segments ................................................................................................... 16 2.3.2.
2.4. Biodegradation of polyurethanes ........................................................................... 19
2.5. Polyurethane scaffolds ........................................................................................... 21
2.6. Injectable polyurethanes ......................................................................................... 23
2.7. Polyurethanes modified by biologically active substances .................................... 24
Nanocellulose - general properties ................................................................... 24 2.7.1.
Preparation of CNC .......................................................................................... 26 2.7.2.
Nanocellulose-polyurethane composites .......................................................... 26 2.7.3.
Nanocellulose in medicine ............................................................................... 27 2.7.4.
3. MAIN AIMS ................................................................................................................ 29
4. EXPERIMENTAL PART ............................................................................................ 30
4.1. Chemicals ............................................................................................................... 30
4.2. Solvent free synthesis of bio-PU ............................................................................ 30
4.3. Synthesis of bio-PU nanocomposites ..................................................................... 30
4.4. Characterization ..................................................................................................... 30
UV-VIS ............................................................................................................ 30 4.4.1.
Attenuated total reflection infrared spectroscopy (ATR-IR) ........................... 30 4.4.2.
Thermogravimetric analysis (TGA) ................................................................. 30 4.4.3.
Differential scanning calorimetry (DSC) ......................................................... 30 4.4.4.
Wide-angle X-ray scattering (WAXS) ............................................................. 31 4.4.5.
Swelling ............................................................................................................ 31 4.4.6.
Hydrolytic degradation ..................................................................................... 31 4.4.7.
Extraction ......................................................................................................... 31 4.4.8.
Tensile measurement ........................................................................................ 32 4.4.9.
Dynamic light scattering .................................................................................. 32 4.4.10.
Dynamic mechanical analysis (DMA) ............................................................. 32 4.4.11.
7
Transmision electron microscopy (TEM) ........................................................ 32 4.4.12.
Cytotoxicity evaluation .................................................................................... 32 4.4.13.
5. RESULTS AND DISCUSSION .................................................................................. 34
6. CONCLUSIONS .......................................................................................................... 35
7. REFERENCES ............................................................................................................. 36
8. LIST OF SHORTCUTS AND SYMBOLS ................................................................. 45
9. APPENDIX .................................................................................................................. 47
10. LIST OF FIGURES ...................................................................................................... 47
11. LIST OF TABLES ....................................................................................................... 47
8
1. INTRODUCTION
Polyurethanes (PUs) form diverse class of elastomers covering broad range of chemical
and physical properties. In addition to industrial applications, their structure can also be
designed to form bioresorbable materials for use in human medicine [1–3]. Important features
of PUs for the use in tissue engineering are tissue-specific biocompatibility, biodegradability,
mechanical flexibility and moderate blood compatibility [4]. PUs have been available in
medicine since the 1960s as bio-inert, hydrolytically stable materials (e.g. ventricle assist
devices or heart valves [5]). Later on, bioresorbable PUs were introduced for various tissue
constructs such as nerve conduits [6], vascular grafts [7], cartilage [8], cancellous bone graft
substitutes [9] or as grafts for small diameter vascular replacement [10].
Moreover, bioresorbable PUs can be tailored to possess a broad range of mechanical
properties by selections and content of soft and hard segments. Soft segments are commonly
oligomeric macrodiols represented by linear polyesters such as poly(-caprolactone),
poly(lactic acid) and poly(glycolic acid) or polyethers like poly(ethylene glycol) and
poly(propylene glycol) having a low glass transition temperature (lower than 25 °C). The hard
segments are provided by the combination of the chain extender and the diisocyanate
component. The nature of hydrogen bonding in the hard segment causes a mutual attraction
leading to a formation of hard and soft segment domains [11]. To eliminate the concern of
aromatic amine by-products, the utilization of aliphatic diisocyanates (as alternatives to
aromatic ones) provide a route to synthetize biodegradable polyurethanes yielding non-toxic
degradation products [3].
Generally, current bioresorbable PUs are prepared in dimethyl sulfoxide or
dimethylformamide (DMF) solvent followed by casting into a mold. This overcomes the
problem with hydrophilic PEG component in PU, which is grabbing easily the water from air
moisture within polymerization. The water reacts fast with isocyanate resulting in bubbles of
carbon dioxide. The same strategy is commonly used for preparation of PU nanocomposites
with cellulose nanocrystals (CNC). PU matrix is dissolved in DMF and CNC are dispersed in
similar solvent. Mixing of those low viscous dispersions is enabling preparation of material
with fine dispersion of nanoparticles. However, avoiding the use of organic solvent is
favorable parameter in the material processing with the potential use in medicine. Therefore,
preparation of either PU itself or PU-CNC nanocomposite in bulk is challenging.
The aim of the work was to develop solvent free procedure eliminating organic solvents
routinely used in PU synthesis. Polymerization conditions yielding compact biodegradable
films were optimized. The experimental system was based on bioresorbable macrodiols from
PEG and PCL and aliphatic 1,6-diisocyanato hexane (HDI), which is claimed as a non-toxic
amine producer during polyurethane degradation [3]. For preparation of nanocomposites was
chosen PU made of only PEG and HDI due to its amorphous nature. Effect of nanofiller on
the matrix is then not shielded by the effects associated with crystallization.
9
A range of bioresorbable PU systems with similar molecular structure [12-17] or
nanocomposites made of PU and CNC [18, 19], has already been described in literature. Thus,
the approach employed in this work to synthesize thermally crosslinked bio-PUs and PU-
CNC nanocomposites without the use of organic solvent is novel and original.
10
2. THEORETICAL PART
2.1. Chemical structure of polyurethanes
Polyurethanes (PU)s are a group of polymers prepared according to the diisocyanate-
polyaddition principle. The name polyurethane is derived from ethylcarbamate, known as
urethane. The reaction of polyisocyanate with polyol leads to the formation of urethane bond,
as illustrated in Figure 1.
OCN R NCO +n
polyaddition
C NH R NH C O
O
CH2 O
O
diisocyanate
n
n
OH O OH
or polyether
polyol
polyester
urethane bond
or polyether
polyester
Figure 1: General formation of polyurethane.
Compared with other functional groups like ether-, ester-, and also urea-groups, the
urethane groups represent often only a minority of the total composition (e.g. 4–6 % in
flexible foams) [20]. Therefore, the properties are not significantly affected by the urethane
groups. Most polyurethanes are composed of at least three components – long chain polyol,
diisocyanate and chain extender.
PU elastomers are characterized by a segment structure (alternating block copolymers) of
the primary chain as showed in Figure 2. The secondary and tertiary structure and
consequently, the morphology of these polyurethanes depend on the chemical composition
and on the length of the segments (blocks). PU elastomers can be microphase separated to
form hard and soft domains due to differences in polarity between the hard (polar) and soft
(nonpolar) segments [11]. The soft segments typically have a glass transition temperature
lower than 0 °C and are therefore rubbery at the room temperature.
11
Figure 2: Modelled primary structure of segmented polyurethane [20].
PUs can be either physically or chemically cross-linked by using the components with
more than two functional groups, e.g., triisocyanate or a multihydroxyl polyol [21].
Hydrogen bonding (Figure 3) between urethane and urea groups in the hard segments of
adjacent polymer chains induces the formation of ordered hard domains, which function as
physical crosslinks that resist flow when stress is applied to the material. To process the
polymer, the physical crosslinks are broken by heating the material above the hard segment
melting transition or by dissolving the material in an aprotic solvent, such as
dimethylformamide (DMF).
Figure 3: Hydrogen bonding between the chains of a polyurethane based on 1,6 – hexane
diisocyanate and 1,4 – butane diol depicted according to [20].
12
2.2. Bioresorbable polyurethanes
It is important to distinguish between biodegradable and bioresorbable materials.
Biodegradability means that the material degrades and moves away from the site of its action
in vivo, but is not necessarily removed from the body. Bioresorbability then means total
elimination of the initial foreign material and degradation of all low molecular weight
compounds with no residual side effects [21].
Important features of PUs for tissue engineering are tissue-specific biocompatibility,
biodegradability, mechanical flexibility and moderate blood compatibility. PUs were initially
used in medicine as biostable materials. Various implantation sites for biostable PUs occurred
in areas such as cardiovascular system (artificial heart [24], vascular grafts [25, 26] and stents
[3, 27, 28]), the middle ear (artificial tympanic membrane [29]), the eye (artificial intraocular
lenses [30]), the digestive tract (stent-like extensions for the oesophagus [31] or the
replacement of biliary ducts [32]).
Applications of bioresorbable polyurethanes appeared later on. As scaffolds they have
already been used for various tissue constructs such as nerve conduits [6], vascular grafts [7],
cartilage [8] or cancellous bone graft substitutes [9]. The usage of polyurethanes in medicine
divided by standards of United States Food and Drug Administration (U.S. FDA) is
summarized in Table 1.
Table 1: Applications of polyurethanes in tissue engineering divided by U.S. FDA standards
[11].
U.S. FDA class I or II: U.S. FDA class III:
• Catheters, drains, medical tubing
• Blood bags
• Transdermal drug delivery patches
• Surgical drapes and gowns
• Transient cardiovascular devices
• Intra-aortic balloon pumps
• Temporary left ventricular assist devices
and biventricular assist devices
• Interim cardiovascular devices
• Ventricular assist devices
• Total artificial hearts used as temporary
measure prior to organ transplantation
(generally intended for less than a month)
• Permanent cardiovascular devices
(e.g. totally implantable ventricles,
artificial hearts, pacemakers and leads,
artificial heart valves)
• Blood conduits and access devices
• Wound dressings and barrier scaffolds
• Contraceptives
Note: Device FDA classification depends on the intended use of the device and also on the indications for
use. In addition, classification is risk based. This means the risk the device poses to the patient and/or the user is
a major factor in the class it is assigned. Class I includes devices with the lowest risk and Class III includes those
with the greatest risk [33].
13
2.3. Feedstocks for synthesis of bioresorbable polyurethanes
The stability of PUs in the body depends on a number of factors including structure and
composition. The synthesis of PUs requires at least two components, the first one is isocyanate
and the second one is a bi- or multi-functional polyol.
The direct reaction of diisocyanate with a long-chain diol usually yields a soft polymer
with low mechanical strength. This property is usually changed by the addition of a chain
extender, which is low molecular weight diol or diamine. Diamine chain extenders yield
poly(urethane urea)s, which have higher hard segment melting temperatures and therefore
harder mechanical properties compared to PUs obtained from diol chain extenders [34]. The
properties of the final PU are primarily dependent on the chemical nature of mentioned three
building blocks and the relative proportions between them used during synthesis.
To increase the reaction rate, urethane catalysts such as tertiary amines or compounds as
organotin and/or elevated temperatures (60–90 °C) may be used. Dibutyltin dilaurate and
stannous octoate prevail in the articles about biodegradable PUs. Stannous octoate is more
often used as catalyst due to its lower cytotoxicity compared to dibutyltin dilaurate [35]. The
structure of both of them is shown in Figure 4.
a)
H3C(H2C)9H2C
O
O
Sn
O CH2(CH2)9CH3
O
CH3
CH3
b)
O
O-
O
O-
Sn2+
Figure 4: Chemical structure of a) dibutyltin dilaurate and b) stannous octoate.
14
Hard segments 2.3.1.
Hard segments are in PU structure represented by isocyanates and chain extenders.
Traditional aromatic diisocyanates are replaced by linear ones to avoid toxic or carcinogenic
degradation products. The structures of polyisocyanates used to synthesize biodegradable PU
biomaterials are showed in Table 2. Important aliphatic isocyanates used are 1,4-
diisocyanatobutane (BDI), hexamethylene diisocyanate (HDI), dicyclohexylmethane
diisocyanate (H12MDI). Lysine-derived polyisocyanates, including lysine methyl (or ethyl)
ester diisocyanate (LDI) and lysine triisocyanate have been extensively investigated, as they
did not show any harmful effects in vivo [36]. As published by Zhang et al. [37] in their tests
on rabbit bone marrow stromal cells, LDI-glucose synthesized polymers induced no
immunogenic reaction and no antibody responses in the host. Their PU did not cause any
accumulation of macrophages (foreign body giant cells) or tissue necrosis around the material.
Table 2: Isocyanates used in polyurethane chemistry for tissue engineering.
Chemical name Structure
1,4-Diisocyanatobutane (BDI)
N C O
NCO
1,6-Diisocyanatehexane
or hexamethylene diisocyanate (HDI)
N C O
NCO
Dicyclohexylmethane diisocyanate
(H12MDI)
N C ONCO
Lysine methyl ester diisocyanate
(LDI)
NCO
OO
N C O
CH3
Reactions of isocyanate are summarized in Figure 5a-e. Water reacts with isocyanates to
form carbamic acid which is an unstable compound and decomposes immediately to an amine
and carbon dioxide gas. The water reaction is used to synthesize PU foams (called in tissue
engineering scaffolds), where the carbon dioxide gas acts as a blowing agent. Isocyanates
react with hydroxyl- and amino-functional molecules by addition to the carbon-nitrogen
15
double bond. Urethane bonds are formed by reaction of isocyanates with hydroxyl-functional
molecules and urea linkages are formed by reaction with amines. The excess of isocyanate is
used to introduce certain degree of crosslinking, which is desirable for improved mechanical
strength, since it leads to the formation of allophanate and biuret branch points [34].
a) With water: amine
R N C O + H2O R NH2C OH
O
R N
H
+ CO2
b) With alcohol: urethane
R N C O OH R'+R
NH
CO
O
R'
c) With amine: urea
R N C O NH2 R'+ R
NH
CO
NH
R'
d) With urethane group: allophanate
R N C OR'
NH
CO
O
R''+
RNH
CON
R'
COO
R''
e) With urea group: biuret
R N C O + R'
NH
CO
NH
R'' RNH
CON
R'
CONH
R''
Figure 5: Reactions of isocyanate with other functional groups.
The relative fractions of the hard and soft segments affect the mechanical properties of
segmented PUs. Hard segment content increases with increasing chain extender to polyol
ratio. This yields harder and stiffer polymers with higher tear strength but lower elongation at
break. Gisselfaelt et al. [38] demonstrated this in PUs made of MDI and 1,3-diaminopropane
as hard segment part and PCL as a soft segment. Reducing the PCL soft segment molecular
weight from 2000 to 530 g·mol-1
increased the hard segment content from 21.9 to 51.4 wt%,
16
resulting in increased stiffness (6 to 45 MPa) and decreased elongation at break (77 to 32 %).
The polymer became harder and less elastic with increasing hard segment content.
Soft segments 2.3.2.
Polyols are viscous liquids with hydroxyl end groups and a polyether, polyester,
polycarbonate, polydimethylsiloxane, or polybutadiene backbone. The majority of
biodegradable PUs are based on polyester or polyether macrodiols. Typical representatives
include polyethyleneglycol (PEG), polypropyleneglycol (PPG), polylactic acid (PLA),
polyglycolic acid (PGA) and their copolymer polylactic-co-glycolic acid (PLGA),
polycaprolactone (PCL) or polyhydroxybutyrate (PHB). Their properties and characterization
have already been reviewed many times [39, 40]. Their major properties are listed in Table 3.
Soft segments affect biostability in the PU materials. Polyester soft segments were found to
be more susceptible to hydrolytic degradation and attack by microorganisms than polyethers
[1].
2.3.2.1. Poly(ethylene glycol)
Poly(ethylene glycol) (PEG), also known as poly(ethylene oxide), poly(oxyethylene) or
polyoxirane, is a hydrophilic, non-ionic polymer. PEG is known for its biocompatibility and
low toxicity and is soluble in water via hydrogen bonding interactions. It has been approved
by the U.S. FDA for internal consumption [41]. It can make a surface highly resistant to
biological fouling and can reduce protein adsorption and resistance to bacterial and animal
cell adhesion. Most PEGs with molecular weight < 1000 g·mol-1
are removed from the body
unaltered only dissolved in the body fluids [42]. It is also apparently not readily recognized by
the immune system. Modifying proteins with PEG have been shown to reduce the
immunogenicity and antigenicity of these proteins and to increase circulation times [43].
PEGs having molecular weight less than 1000 g·mol-1
are viscous, colorless liquids. Higher
molecular weight PEGs are waxy, white solids. The melting point of the solid is proportional
to molecular weight, approaching a plateau about 67 °C. The molecular weights commonly
used in biomedical applications range from a few hundreds to 20 000 g·mol-1
[42].
Block copolymer of poly(ethylene-propylene-ethylene glycol) is sold under the trademark
Pluronic®[45] and is often used in synthesis of bioresorbable PUs. Gorna and Gogolewski
[46] synthesized linear biodegradable polyurethanes with varying ratios of the hydrophilic-to-
hydrophobic segment. The hydrophilic segment was based on Pluronic® and the hydrophobic
segment was based on PCL diol. The polymers absorbed up 3.9 % of water depending on the
chemical composition. Tensile strength, Young's modulus and elongation at break of the
polymers were in the range of 11–46 MPa, 4.5–91 MPa and 370–960 %. Degradation in vitro
caused 2 % of mass loss in 48 weeks. The extent of degradation was dependent on the
polymer composition and the hydrophilic segment content.
17
Table 3: Physical properties of the biodegradable polymers used as polyols in formation of polyurethane.
Synthetic polymer Thermal and mechanical properties Degradation properties Applications Ref.
Tm (°C) Tg (°C) Elastic
modulus (GPa)
Time
(months) Products
PEG - Upper limit
at 67 °C - Immediately
Dissolves in body
fluids
Laxatives, bowel preparation
before surgery, carrier for active
ingredients
42
PCL 58-63 -60 0.4 >24 Caproic acid Sutures, dental orthopaedic
implants 47
PGA 225-230 35-40 5-7 3-4 Glycolic acid Suture anchors, meniscus repair,
medical devices, drug delivery 50
PLA 170 56 8.5 12-18 L-lactic acid Fracture fixation, suture anchors,
meniscus repair 40, 50
PLGA (50/50) Amorphous - 2.0 1-2 weeks D,L-lactic acid and
glycolic acid - 43
PLGA (85/15) Amorphous - 2.0 5-6 weeks D,L-lactic acid and
glycolic acid
Interference screws, suture
anchors, ACL reconstruction 43
PHB 177 2 3.5 >17 D-3-hydroxybutyric
acid Screws 55, 57
18
2.3.2.2. Poly(-caprolactone)
Poly(-caprolactone) (PCL) is a semicrystalline, bioresorbable polymer belonging to the
aliphatic polyesters. PCL has melting point ranging between 59 and 64 °C and a glass
transition temperature of about −60 °C. PCL can be prepared either by ring-opening
polymerization of -caprolactone using a variety of anionic, cationic and co-ordination
catalysts or via free radical ring-opening polymerization of 2-methylene-1-3-dioxepane [47].
It is regarded as a soft and hard tissue compatible bioresorbable material and has been used in
biomedicine e.g. as a scaffold, sutures or fixation devices [48].
The PCL homopolymer has a total degradation of 2–4 years depending of the starting
molecular weight used for the implant. Its degradation pathway undergoes a two-stage
degradation process. First, the non-enzymatic hydrolytic cleavage of ester groups, and second,
when the polymer has lower molecular weight (less than 3000) it undergoes intracellular
degradation as evidenced by observation of PCL fragments uptake in phagosomes of
macrophages and giant cells and within fibroblasts [47]. Local decrease of pH during the first
stage has been observed same as in other polyesters. The combination of PCL with other
materials such as bioceramics can suppress the local decrease of pH during degradation [49].
2.3.2.3. Polyglycolic acid
Polyglycolic acid (PGA) is a semicrystalline biodegradable polymer known to lose its
strength in 1–2 months when hydrolyzed and losses mass within 6–12 months. PGA sutures
(Dexon®) have been commercially available since 1970 [50]. The hydrolytic degradation in
vivo may take place via nonspecific enzymes such as esterases and carboxyl peptidases that
produce glycolic acid monomers. Glycolic acid monomers are then converted enzymatically
either into glycine, which can be used in protein synthesis, or into pyruvate, that enters the
tricarboxylic acid cycle and yielding energy, CO2 and water [51]. Glycolic acid is partially
excreted in urine. Hydrolysis is affected by the initial molecular weight, surface area/weight
ratio, porosity and monomer concentration, geometric isomerism, conformation and
crystallinity. Increased pH accelerates PGA degradation.
2.3.2.4. Polylactic acid (PLA)
Polylactic acid (PLA) is a thermoplastic, high strength, high modulus polymer, which
belongs to the family of aliphatic polyesters. Poly(L-lactide) is a crystalline polymer
(approximately of 37 % crystallinity) while poly(D,L-lactide) is amorphous. Poly(L-lactide)
has a glass transition temperature between 60–65 °C and a melting temperature approximately
of 175 °C, while poly(D,L-lactide) has glass transition temperature between 44–55 °C [40].
PLA undergoes scission of chains in the host body to monomeric units of lactic acid, which is
a natural intermediate in carbohydrate metabolism. These characteristics make this polymer
suitable for use as resorbable sutures, carries for the controlled release of drugs, implants for
19
orthopaedic surgery or blood vessels, which finally can be replaced by living tissues. The
attraction of PLA as a biodegradable material is its ready availability from renewable
resources such as corn starch (in the U. S.) and sugarcanes (rest of world). The PLA life cycle
starts with corn starch. The plants are first milled to separate the starch, which is converted to
lactic acid utilizing fermentation and a series of purification steps. It is fully compostable. It
can be converted back to monomer and oligomer by enzymatic degradation, or it can be
degraded into water, carbon dioxide and organic materials [52].
2.3.2.5. Poly(lactic-co-glycolic acid)
Poly (lactic-co-glycolic acid) (PLGA) copolymer is built from the units of poly(lactic acid)
and poly(glycolic acid). The rate of polymer degradation may affect many cellular processes,
including cell growth, tissue regeneration, and host response. PLGA has been shown to
undergo bulk erosion through hydrolysis of the ester bonds and the rate of degradation
depends on a variety of parameters including the PLA/PGA ratio, molecular weight, and the
shape and structure of the matrix. The degradation products are endogenous compounds
(lactic and glycolic acid) and are nontoxic. The rate of degradation can be controlled by the
ratio of PLA and PGA. The higher the content of glycolide units the faster is the degradation
[53]. PLGA has good cell adhesion and proliferation. Various studies have been performed so
far using micro- and nano-fabrication techniques to form three-dimensional scaffolds based
on PLGA [43]. Another application of biodegradable PLGA is for use in guided tissue
regeneration by providing a permeable material for space preservation or in drug delivery
systems [54].
2.3.2.6. Polyhydroxybutyrate
Poly-3-hydroxybutyrate (PHB) belongs to a group of polyesters called
polyhydroxyalkanoates (PHAs). PHAs are biodegradable plastics produced by
microorganisms. They have mechanical properties similar to some common synthetic
thermoplastics and are environmentally degradable, where the products of decomposition are
CO2 and H2O [55]. PHB has mechanical properties quite similar to commercial polypropylene
with Young’s modulus 3.5 GPa, tensile strength of 40 MPa and melting point at 179 °C.
Moreover, product of PHB degradation (3-hydroxybutyric acid) belongs to short-chain fatty
acids and reveals antibacterial activity [56].
2.4. Biodegradation of polyurethanes
PUs were firstly used as a biostable materials in medicine and since that time have been
applied in a range of medical devices [58, 59] In late 1980s a failure on functions in the
applications such as the pacemaker leads and breast implant coatings containing PUs was
found [3]. Aromatic diisocyanates (e.g. methylene diphenyl diisocyanate (MDI) and toluene
diisocyanate) have been recognized as potentially carcinogenic by-products of degradation.
20
To solve this problem, aliphatic diisocyanates such as 1,4-diisocyanatobutane (butyl
diisocyanate, BDI) and lysine diisocyanates (LDI) started to be used instead of aromatic ones.
These polymers exhibit better blood and tissue compatibility both, in vitro and in vivo tests, in
comparison to materials where aromatic reactants were used [39].
PU elastomers with hydrophilic soft segments have been reported to exhibit increased
water uptake, which has also been suggested to increase the degradation rate [38]. Mechanism
of hydrolytic degradation that has been suggested by Hafeman et. al [63] is shown in Figure 6.
Above mentioned phenomenons can be used to design PU materials with targeted physical
properties.
Figure 6: Schematic showing hypothesized mechanisms of PU scaffold degradation and
soluble breakdown products. (A) Hydrolysis of ester bonds to yield -hydroxy acids (dotted
line). (B) Abstraction of the -hydrogen atom in the R group adjacent to the urethane bonds
by reaction with hydroxyl radicals to yield lysine or ethanolamine and carbon dioxide. (C)
Abstraction of the -hydrogen atom in lysine by reaction with hydroxyl radicals to yield
multiple degradation products, including NH4, -ketoacids, oximes, CO2, and carboxylic
acids (degradation products not shown to simplify the diagram) [63].
The stability of segmented (block) PUs in the body depends not only on the composition
and structure. It also depends on the presence of stabilizing additives, the chosen implant site
and the sample’s purity and processing history. Fabrication technique, handling, residual
solvent, sterilization method and device design are also important factors [23]. The common
degradation pathway of poly(ester urethanes) is the hydrolysis of ester linkages. Through this
way the degradation products are hydroxy acids as well as urethane and urea fragments with
terminal acid groups. The hydrolysis of urethane and urea linkages to amine has been also
21
reported [60]. Hydrolysis of ester bonds is taking control of the degradation rate. The
formation of acidic degradation products is known to autocatalyze polymer degradation and
may cause in vivo inflammation [60].
By incorporating a chain extender that is recognizable by an enzyme, PU elastomers with
degradable hard segments have been also synthesized [61]. Guan et al. [62] synthesized PUs
from BDI, PCL-PEG-PCL macrodiol triblock copolymers and the peptide Alanine-Alanine-
Lysine (AAL) as a chain extender. The hard segment of this polymer was designed to be
enzymatically degradable. Polymers synthesized from the AAL chain extender degraded
faster than those synthesized from tetramethylenediamine (putrescin®) chain extender and the
degradation products were observed to be non-cytotoxic. Further, materials synthesized from
the AAL chain extender displayed mechanical properties comparable to those synthesized
from usually used chain extender putrescin®. Obtained tensile strengths ranged from 15 to
28 MPa, breaking strains 670–890 %.
2.5. Polyurethane scaffolds
Many techniques/methods are employed to process biomaterials into scaffolds. The reason
why to make scaffolds rather porous than solid, is that cell colonization, in-growth and
proliferation can proceed easier. Porosity leads to a weaker structure and it can also accelerate
degradation due to the larger surface area. Having a porous structure allows nutrients to flow
into the cells in the inner regions of the scaffold and for waste to be removed. The ideal pore
size and pore morphology of the scaffold depends on the intended purpose of the scaffold. It
has been claimed that the optimal pore size is between five and ten times the diameter of the
cell i.e. 100–300 μm [65].
Techniques used for processing of scaffolds have been reviewed many times [65, 66].
Because the dissertation thesis is not directly focused on preparation of scaffold, there is just a
list of them. Techniques involve solvent casting and particulate leaching, gas foaming, non-
woven fibers, fiber knitting and phase separation/emulsion freeze drying. Emerging
fabrication techniques under development include solid free forming techniques, including 3D
printing and fused deposition. Characteristics differentiating the techniques are the use of
solvents, heat, pressure, or pore creating agents.
The scaffold should be non-immunogenic, non-toxic, biocompatible, biodegradable and
easily manufactured. The cells survival and growth is affected by macro and micro-structural
properties. Scaffold should possess an interconnected and spread porosity (usually exceeding
90 %) with a highly porous surface and microstructure. This would allow in vitro cell
adhesion and would provide the necessary space for neovascularization in vivo. Pore
interconnectivity directly influences the diffusion of physiological nutrients and gases to cells
as well as the removal of metabolic waste and by-products from cells.
22
Grenier et al. [67] examined two medical grade polyurethanes for their potential use as
scaffold for vascular tissue engineering applications. Those two polyurethanes were
commercially available Bionate 55D (polycarbonate-urethane – thermoplastic elastomer
formed from hydroxyl terminated polycarbonate, an aromatic diisocyanate, and a low
molecular weight glycol as a chain extender [68]) and Elasthane 75D (polyether-urethane –
made of poly(tetramethylene oxide) and an aromatic diisocyanate and a low molecular weight
glycol chain extender [68]). Scaffolds were prepared via pressure differential/particulate
leaching technique, which improved the pore interconnectivity of the scaffolds. Two types of
porogen were used for fabrication of the scaffolds, NH4Cl particles, and paraffin spheres. In
the preparation procedure, they firstly dissolve the polyurethane in dimethylformamide. Then
the cylindrical mold was filled with porogen and using the pressure dissolved polymer was
pushed into the mold. After a certain time the mold was removed from the apparatus and
placed in a fume hood at ambient temperature for 2 days to evaporate the solvent. The
polymer matrix was then pushed out of the mold and immersed in deionized water for NH4Cl
particles or hexane for paraffin spheres, to leach out the porogen.
Sharifpoor et al. [69] synthesized degradable polar hydrophobic ionic polyurethane (D-
PHI) porous scaffold. D-PHI scaffold properties were adjusted through the introduction of a
lysine-based cross-linker. Increasing lysine-based cross-linker concentration resulted in an
increase of the elastic modulus (from 0.5 to 21 MPa), a decrease of the elongation-at-yield
(from 45 to 5 %) and a reduction of scaffold swelling (from 170 to 100 %). Based on a
preliminary study with vascular smooth muscle cells, D-PHI scaffolds demonstrated the
ability to support cell adhesion and growth during 2 weeks of culture, suggesting their
potential suitability for longer term vascular tissue engineering. The versatility of the D-PHI
properties may allow for the tailoring of cell-material interaction and ultimately functional
tissue regeneration.
23
2.6. Injectable polyurethanes
Synthetic polymer systems as injectable liquids, gels or pastes have the advantage for the
surgical operations, because they minimalize the invasive procedures. The way how to deliver
the gel to the proper place in the body is for example arthroscopic delivery. Injectable
polymer systems could help in complex fracture and bone defect repair, where substantial
amount of bone is lost due to trauma or disease. Further, they can be used to stabilize and
reinforce the fixation of implants such as plates and screws, particularly in patients with
osteoporotic bone [70]. Several biodegradable injectable hydrogel systems made of
polyurethanes have been suggested as described below.
Thermoresponsive polymers showing enhanced rheological properties were published by
Cohn et al. [71]. PUs were synthesized from PEO-PPO-PEO triblock macrodiols end capped
with lactid acid or -caprolactone oligoesters and chain extended with HDI. The length and
composition of the ester blocks affected the viscosity response with temperature. Gels with
viscosities up to 193 kPa∙s were prepared. Degradation times ranging from days to months
were achieved by varying the composition of the backbone. The improved rheological
properties of these materials were suggested as potential injectable drug delivery system.
Guelcher et al. [72] introduced two-component injectable polyurethane hydrogels as useful
biodegradable material for fracture healing. PU network consisted of reactive liquid molding
low-viscosity prepolymers derived from lysine polyisocyanates and poly(caprolactone-co-
D,L-lactide-co-glycolide) triols. The values of Young’s modulus ranged from 1.20
to 1.43 GPa, and the compressive yield strength varied from 82 to 111 MPa, which is
comparable to the strength of poly(methyl methacrylate) bone cements. They studied the
biodegradation in vitro, where the materials decomposed to non-cytotoxic products. As well,
the material supported the attachment and proliferation of human osteoblasts.
An injectable, two component LDI-based polyurethane system with commercial name
PolyNovo® [73] was developed for orthopaedic applications with curing in situ. This self-
setting system can be inserted to the body arthroscopically in liquid form. It polymerizes at
physiological temperature in situ to provide appropriate bonding strength and mechanical
support comparable to or superior to widely used bone cements. This material has also been
shown to promote favorable cell adhesion and proliferation [74].
Although quite a lot of articles dealing with injectable PU concept have been published,
the injectable PUs have still basic problems that should be overcome. The first one is the
toxicity and ultimate fate of reactive components that are not incorporated in the final cured
product [75]. The nature of the reaction or reactivity of the NCO-terminated prepolymer is
exothermic, which may destroy the surrounding tissue. And finally, diffusion of water from
the wound bed into reactive PU can result in over-expansion of the formed scaffold or in
forming large voids in the scaffold.
24
2.7. Polyurethanes modified by biologically active substances
Each group, either natural or synthetic polymers, has its advantages and disadvantages for
the use in tissue engineering. Synthetic polymer-derived materials lack cell recognition
signals and the surfaces are hydrophobic, which affect the seeding of cells. Naturally derived
polymers such as collagen have the advantage of good cell interaction and hydrophilicity, but
generally lack adequate mechanical properties and batch-to-batch consistency [76]. To
overcome the drawback of the synthetic materials, naturally occurring polymers have been
used to modify the synthetic materials. Polyurethanes have been used for a long time as
potential candidates in various fields of biomedicine. Thus, different natural polymers were
already used for their modification, such as collagen [77, 78], elastin, [79-82], hyaluronic acid
[83-86], silk fibroin [87, 88], chitosan [89-91], bioactive glasses [92-94] or nanocellulose.
The last mentioned natural polymer was used in the presented work in the form of cellulose
nanocrystals, and therefore is reviewed more into detail in the following section.
Nanocellulose - general properties 2.7.1.
As natural nanoscaled material, nanocellulose provides unique properties different from
other nanomaterials, including special morphology and geometrical dimensions, crystallinity,
liquid crystalline behavior, alignment and orientation, mechanical reinforcement,
biocompatibility, biodegradability and lack of toxicity [95]. Nanocellulose can be divided into
three different types such as cellulose nanocrystals (CNC), cellulose nanofibrils (CNF) and
bacterial cellulose (BC) also called microbial cellulose [96].
Cellulose nanocrystals with other designations such as nanocrystalline cellulose, cellulose
(nano) whiskers or rod-like cellulose microcrystals are nanoparticles with diameter in a range
of 5–30 nm, and length between 100–500 nm derived from plant cellulose, or length of 100
nm to several micrometers from tunicate and algae celluloses. Acid hydrolysis is commonly
performed for the extraction of CNC through the removal of amorphous regions and
preservation of highly-crystalline structure. Scheme of their structure is shown in Figure 7.
25
O
C1
C2
C3
C4C5
O
C4OHOH
C6OH
C5
C3
C2
C1
O
O
C6OH
OHOH
n
1 4
cellulose chainsdiordered region
crystalline regions
cellulose nanocrystals
100 nm
a)
b)
c)
Figure 7: Schematics of (a) single cellulose chain repeat unit, showing the directionality of
the 1 -> 4 linkage and intrachain hydrogen bonding (dotted line), (b) idealized cellulose
microfibril showing one of the suggested configurations of the crystalline and amorphous
regions, and (c) cellulose nanocrystals after acid hydrolysis dissolved the disordered regions.
Figure depicted according to [97].
Different from rigid CNC, cellulose nanofibrils (CNF) consist of both individual and
aggregated nanofibrils made of alternating crystalline and amorphous cellulose domains,
which attributes to the morphology of CNF with soft and long chains. Due to the
entanglement of long cellulosic chains, it is not so easy to determine the length of CNF with
microscopic techniques. Therefore, only the information of fibril width for CNF is provided
in the studies, which varies from 10 to 100 nm depending on the source of cellulose,
defibrillation process and pretreatment. Regarding the preparation of CNF, mechanically
induced destruction processes are mainly applied, which involves high-pressure
homogenization and grinding before or after chemical or enzymatic treatment. Multiple
mechanical shearing effectively delaminate individual microfibrils from cellulosic fibers.
Contrary to CNC and CNF, bacterial cellulose (BC) is produced by biosynthesis typically
by bacteria such as Acetobacter xylinum [98]. The major advantage is that the product is in a
pure form and doesn’t require removing of unwanted impurities or contaminants such as
lignin, pectin and hemicellulose. During the biosynthesis of BC, the glucose chains are
produced inside the bacterial body and extruded out through pores present on the cell
envelope. With the combination of glucose chains, microfibrils are formed and further
26
aggregate as ribbons (nanofibers). These ribbons subsequently generate a web-shaped
network structure with cellulosic fibers (BC), which has a diameter of 20–100 nm with
different types of nanofiber networks.
Preparation of CNC 2.7.2.
Considerable academic and industrial research effort have been devoted to developing
cellulose nanocrystals and nanofibers. The reasons for this are e.g. that cellulose is one of the
most abundant polymeric material in nature, is available everywhere, can be obtained from
different resources, is biodegradable and can be modified to generate whiskers of high aspect
ratio, high modulus and low density [99]. To obtain these nanofibers, cellulose must be
separated from other components of the plant and exposed to different treatments that
decompose its hierarchical structure down to nanoparticles. One frequently reported method is
the acidic hydrolysis of cellulose, which produces cellulose nanocrystals with diameter of 5–
20 nm and length ranging from 100 nm to approximately 500 nm, depending on the cellulose
source and the preparation process [100, 101] Although the high OH concentration on the
surface of the crystals suggests high attraction between them, the use of sulfuric acid during
hydrolysis is known to leave sulfate groups on the crystal surface, which produce repulsion
and help to produce stable suspensions in aqueous media and highly polar solvents [102].
The surface hydroxyl groups on the cellulose can be reacted to produce surface-modified
cellulose crystals [101]. Coreaction with a polymerizable mixture has also been reported, so
that the nanofibers became covalently attached to the polymer [102].
Nanocellulose-polyurethane composites 2.7.3.
A couple of systems suggesting PU-CNC composites have been introduced in the
literature. Those are usually using CNC dispersed in N,N-dimethylformamide (DMF) and PU
matrix dispersed in the same solvent. PU nanocomposite films are then prepared by casting
the mixture of those two components into the mold and after removing the solvent the films
are obtained with thickness ranging from 0.1 mm to 0.5 mm.
Pei et al. [18] observed a strong reinforcement effect by adding CNC into PU matrix made
of 4,4’-diphenylmethane diisocyanate (MDI) and 1.4-butanediol as hard segment and
poly(tetramethylene glycol) as soft segment. CNC were grafted by MDI before the synthesis
proceeded. With only 1 wt.% of cellulose nanocrystals incorporated, an 8-fold increase in
tensile strength and 1.3-fold increase in strain-to-failure were achieved. It is unusual to
improve the modulus while at the same time significantly improve the strength and toughness.
Such enhancement of mechanical properties was explained by CNC reinforcement in the hard
domains and also by increased effective cross-link density of the elastomer network due to
CNC-PU molecular interaction. Preferential reinforcement of CNC in the hard microdomains
rather than in the soft segments of PU avoided the undesired stiffening of the soft domain and
maintained the large strain-to-failure.
27
The similar strategy with different PU system performed Rueda et al. [19]. The surface of
CNC was grafted by hexamethylene diisocyanate (HDI) and the PU nanocomposite was
processed via similar solvent casting procedure as Pei et al. Successful grafting was
confirmed by solid state NMR and it was claimed that half of the C6-OH groups in glucose
units were grafted by HDI. Comparison between nanocomposite material with unmodified
CNC and HDI modified CNC, both in 1.5 wt %, showed more tough material when used
unmodified CNC without impeding the extension of the soft segment and provoked a
positively impact on the mechanical properties. The use of HDI modified CNC resulted in
preferential interaction of HDI anchored chains with hard segment which provided fairly good
Young’s modulus however showed decrease to more than half of values in ductility and strain
at break. The same group published later the cell evaluation tests of those PU nanocomposites
[103]. Cells seeded on top of the materials showed that L-929 fibroblasts colonized the
materials surface giving rise to good substrates for cell adhesion and proliferation.
Saralegi et al. [104] used renewable resources, such as castor oil based highly crystalline
polyol and corn-sugar based chain extender. CNC PU nanocomposites were prepared where
PU matrix was partially crystalline. They observed increase in Young’s modulus and yield
strength by increasing CNC content and at the same time decrease in strain at break and
tensile strength. Such behavior of the material was explained by CNCs influencing the
mobility of amorphous segments. However due to high crystalline nature of both soft and
hard segments, that also served as reinforcing agents, the effect of the addition of CNC on the
final properties was less pronounced.
Nanocellulose in medicine 2.7.4.
One of the targeting outcomes of that dissertation work is utilizing the materials for further
use in medicine. Nanocellulose as a unique natural material has gained much attention for its
use in medicine. The studies of blood vessel replacement are the most developed reaching the
stage of clinical research as various systems have been already tested on animals [105].
Regarding other studies on nanocellulose e.g. for soft tissue replacement and regeneration,
most reports are still in the fundamental stage, and mainly focus on the comparison of
different properties between nanocellulose-based materials and real organs. Due to the most
developed section of vessel replacement studies some of them are described more into detail
below.
One of the most common treatments of cardiovascular disease is the graft surgery, which is
performed to supply blood to the damaged tissue (e.g. heart) with a suitable blood vessel
replacement. Biosynthesized BC has already been processed by researchers to form artificial
vascular substitutes. The team of Dieter Klemm (University Jena and Polymer Jena,
Germany) was the first research organization to investigate and apply artificial vascular
substitute obtained from BC. They have evolved a clinical product named BActerial
SYnthesized Cellulose (BASYC®) with high mechanical strength in wet state, enormous
28
water retention property and low roughness of inner tube surface [105]. In comparison with
conventional synthetic vascular graft materials, e.g. polyester (Dacron®) or ePTFE,
biosynthetic BC tubes can be suitable for small diameter (<4 mm) vascular conduits. It has
been reported that BASYC from BC has been successfully used as the artificial blood vessel
in rats and pigs for microsurgery [106].
Different from BC biosynthetic procedure, it is impossible to directly fabricate the tubes
from CNC and CNF. Therefore, the development of CNC or CNF-based blood vessel
replacement commonly includes the use of a matrix material. Novel biomaterials from
polyurethane reinforced with CNF have been reported to be potentially used as vascular
replacement [107]. The presence of CNF in polyurethane improved the elastic properties of
the material, coupled with low thrombogenicity and improved physical and mechanical
properties. CNF/polyurethane biomaterials, with a wall thickness of 0.7–1.0 mm, were
applied as vascular prostheses between the brachiocephalic trunk and the right common
carotid artery in a 26-year-old male patient with multiple endocrine neoplasia. No further
effect of this CNF/polyurethane biomaterial in clinical study was reported yet.
29
3. MAIN AIMS
The dissertation thesis is focused on synthesis of biodegradable polyurethane elastomers
and their modification by biologically active cellulose nanocrystals. Firstly, the PU material
was synthesized with tailorable mechanical and biodegradable properties. For preparation of
PU elastomers were chosen hydrophilic poly(ethylene glycol) and hydrophobic
poly(caprolactone) as diols, aliphatic hexamethylene diisocyanate as hard segment and
stannous octoate as catalyst. Characterization of synthesized materials was evaluated from the
viewpoint of changing the diol ratios (hydrophilic vs. hydrophobic nature) and changing the
crosslink density (ratio between NCO/OH groups). Both parameters possess ability to tailor
the physico-chemical properties for potential application in medicine.
Further modification by cellulose nanocrystals (CNCs) was performed to study the
nanofiller influence on material behavior and also for possible improvement of the
biocompatibility. CNCs were used either untreated or surface modified by polyethylene
glycol.
Proposed dissertation thesis can be summarized in following highlights:
solvent free synthesis of biodegradable PU material
interpretation of PU behavior with different PEG/PCL ratio
tailoring the properties of material by changing crosslink density (NCO/OH ratio)
solvent free synthesis of PU nanocomposite
structural investigation of the nanocomposite’s behavior
30
4. EXPERIMENTAL PART
4.1. Chemicals
Text is not available in the public version.
4.2. Solvent free synthesis of bio-PU
Text is not available in the public version.
4.3. Synthesis of bio-PU nanocomposites
Text is not available in the public version.
4.4. Characterization
UV-VIS 4.4.1.
Absorbance spectra were obtained on Jasco V-630 UV-VIS spectrophotometer in the range
of wavelengths from 240 to 1100 nm with the resolution of 0.5 nm.
Attenuated total reflection infrared spectroscopy (ATR-IR) 4.4.2.
Attenuated total reflection infrared spectroscopy (ATR-IR) was conducted on Fourier-
transform infrared spectrometer (Bruker Tensor 27, USA) equipped with the germanium
crystal for ATR over the spectral range from 4000 to 600 cm-1
at the resolution of 4 cm-1
and
32 scans.
Thermogravimetric analysis (TGA) 4.4.3.
Thermal decompositions were investigated by performing TGA (TA Q500, USA) with
sample mass in the range of 5–10 mg. All measurements were carried out in the nitrogen
atmosphere with the nitrogen flow rate of 100 ml·min-1
. Heating rate was set as 10 °C·min-1
starting from 25 °C up to 750 °C.
Differential scanning calorimetry (DSC) 4.4.4.
The calorimetric measurement was carried out utilizing DSC (Netzsch 204 F1, Germany)
in the nitrogen atmosphere. Each sample (5–10 mg) was cooled with the rate of 5 °C·min-1
and heated with heating rate of 10 °C min-1
. Raw data were processed using the NETZSCH
Proteus® Software to obtain glass transition temperature (Tg), melting temperature (Tm) and
melting enthalpy (Hm). The crystallinity (c) was calculated from the first DSC heating run
according to the following equation:
c (%) = (∆𝐻𝑚 × 100)/(𝑤 × ∆𝐻𝑚0 ) (1)
where ∆𝐻𝑚 is heat of fusion of the bio-PU sample, ∆𝐻𝑚0 is heat of fusion of 100 % crystalline
PCL (135.44 J·g-1
) [108] and w is the PCL’s weight fraction in the bio-PU sample.
31
Wide-angle X-ray scattering (WAXS) 4.4.5.
Wide-angle X-ray scattering (WAXS) was performed on the diffractometer (Rigaku
MiniFlex 600, Japan) with Cu K radiation beam ( = 1.5406 Å) operated at 40 kV and 15
mA in the scattering angle (2 range from 5° to 35° using the scan rate of 2° min-1
with step
of 0.02°. As a reference material, WAXS of a commercial PCL (Mn = 80 000 g mol-1
) was
also measured.
Swelling 4.4.6.
Swelling ratio of the 1 cm × 1 cm × 1 mm specimens cut out of the bio-PU films was
measured in ultrapure water (UPW) at laboratory temperature. The samples weight was
measured within the first hour every ten minutes, within the second hour every 20 minutes
and then every hour for the next 4 hours, with the last measurement performed after 1 day
since the beginning of the swelling experiment. Water uptake was calculated according to the
formula (2):
Water uptake (%) = (𝑤𝑠 − 𝑤𝑑) 𝑤𝑑⁄ × 100 (2)
where ws is the weight of swollen sample at the given time and wd is the weight of dry sample.
Each sample was measured 3 times.
Hydrolytic degradation 4.4.7.
Hydrolytic degradation tests employing the swelled samples were carried out in an
incubator at 37 °C in UPW for 365 days. Each measurement is an average of 3 specimens.
The specimens were removed at the given time from vials with distilled water, wiped using
filter paper and weighed to determine the weight loss. The UPW was changed every two
weeks. Mass loss was calculated according to formula (3):
Mass loss (%) = (𝑤0 − 𝑤𝑡) 𝑤0⁄ × 100 (3)
where w0 is the weight of swollen sample after 24 hours in water and wt is the weight of the
sample in the given time.
Extraction 4.4.8.
Extractions were performed using Soxhlet apparatus. Small pieces of dry films (1×1 cm
and 1 mm thick) were used for the extractions. The extractions were performed for 8 hours
with volume 200 ml of solvent, extracted samples were dried in the vacuum oven to the
constant weight. Extracts from acetone with dissolved residues were evaporated followed by
drying in the vacuum oven to be evaluated by ATR-IR. Bio-PU films with NCO/OH ratio 1.2
were extracted in acetone and in N,N-dimethylformamide (DMF). Both extractions either in
acetone or DMF were carried out via the same procedure.
32
Tensile measurement 4.4.9.
Tensile tests were performed employing the tensile tester (Zwick Roell Z010, Germany)
according to the ISO 527 standard. Dog-bone specimens obtained according to ISO 527-2/5B
were 35 mm in length, 2 mm in width in the middle part, 1 mm thick and gauge length was
10 mm. The 500 N load cell was used for the measurement with the cross-head speed of
10 mm·min-1
corresponding to the 100 %·min-1
deformation rate. The 0.1 N preload was used
and all tensile measurements were performed at the laboratory temperature. Each sample was
measured at least 5 times and data were averaged to get the standard deviation.
Dynamic light scattering 4.4.10.
Dynamic light scattering (DLS, DynaPro Nanostar, Wyatt) using digital autocorrelator and
laser with wavelength of 658 nm was performed to qualitatively prove the size distribution of
nanoparticles. DLS was measured from a dilute suspension of either CNC or CNC-PEG
nanoparticles.
Dynamic mechanical analysis (DMA) 4.4.11.
Dynamic mechanical analysis (DMA) was measured on a TA Instruments Q800 in tension
mode with frequency 1 Hz and the temperature scanning rate 3 °C·min-1
. Samples with
thickness of 1 mm were cut to stripes with dimensions of 4 mm width and 15 mm length.
Transmision electron microscopy (TEM) 4.4.12.
Transmission electron microscopy (TEM) was carried out using Tecnai G2 Spirit Twin 12
microscope (FEI, Czech Republic) Thin sections (∼60 nm) were cut at –80 °C by an
ultramicrotome (Ultracut UCT, Leica, Germany) under cryogenic conditions (the sample and
the knife temperatures were –85 °C and –50 °C, respectively). The ultrathin sections were
transferred to a microscopic grid and observed in the bright field mode at the acceleration
voltage of 120 kV.
Cytotoxicity evaluation 4.4.13.
Chosen samples were tested on cytotoxicity by seeding mesenchymal stem cells (MSCs)
on bio-PU films and evaluation of spreading and proliferation after 24, 72 and 168 hours.
Chosen bio-PU or PU nanocomposite films with a diameter of 6 mm and a thickness of
1 mm were sterilized using combination of UV for 15 min and ethanol. Before seeding with
MSCs, the films were incubated in culture medium (Dulbecco's Modified Eagle's medium
with L-glutamine, 10% fetal bovine serum, 1% penicillin/streptomycin) overnight at 37 °C.
Cells were seeded on PU films at a density of 3·104 cells·cm
-2 in 48-well plate. PU films with
seeded MSCs were cultivated at 37 °C and 5% CO2 atmosphere and cell viability was
acquired in 24, 72 and 168 hours after seeding. The medium was changed every 3 days.
Fluorescence microscopy and live/dead staining (calcein-AM/ propidium iodide) were used to
33
determine the cell viabillity. The mix of calcein-AM (2 μM) and propidium (1,5 μM) in
phosphate buffered saline (PBS) was added to films containing seeded cells and incubated for
15 min at 37 °C and 5% CO2 for live/dead cell detection. The cells were visualized using a
Zeiss Axio Imager 2 microscope.
34
5. RESULTS AND DISCUSSION
Text is not available in the public version.
35
6. CONCLUSIONS
New solvent free method was successfully used for the synthesis of biodegradable
polyurethanes (bio-PUs) and also for in situ synthesis of polyurethane nanocomposites filled
by either unmodified or PEG grafted cellulose nanocrystals.
Different diol ratios between PEG/PCL resulted in materials from brittle through elastic up
to tough elastomeric films. The structural analysis revealed that crystalline PCL domains
formed below the PEG/PCL weight ratio of 30/70 are responsible for abrupt increase in the
elastic modulus, stress and deformation at break by order of magnitude. These domains
reinforce the PU network and their small size allows for more uniform distribution of strain
resulting in enhanced toughness. Extraction studies confirmed crosslinked structure of bio-
PUs although only bifunctional feedstocks were used.
Variation between NCO/OH groups (isocyanate ratio) confirmed increase in crosslink
density with increasing isocyanate ratio. Higher crosslink density reduced the ability of bio-
PUs to absorb water and followly prolonged the time of hydrolytic degradation. Systems with
very low crosslinking tend to be weak and flexible, whereas polymers with high degrees of
crosslinking were more rigid.
Polyurethane nanocomposites filled by either unmodified or PEG grafted CNCs were
successfully synthesized by evolved solvent free method. Presence of nanofiller promoted
significant effect on stiffness explained by restricted motion of the rubbery PU matrix due to
the presence of stiff and rod-like cellulose nanocrystals.
Five representatives from all the samples were chosen for preliminary tests of human
mesenchymal stem cells (MSCs) response to the material. MSCs didn’t adhere on PUs with
high PEG content although the material was modified by low amount of biologically active
CNCs. PUs with high abundance of PCL allowed cell adhesion on the surface, but promoted
with timescale of one week detachment of the MSCs from the surface.
Tailored biomaterials with adjustable functional properties are desirable for many
applications ranging from drug delivery to regenerative medicine. Presented variation of PU
materials possess great range of mechanical (toughness, flexibility) and physical (swelling,
degradation) properties. Low adherence of cells to the bio-PU surface predestines the material
for the use as an artificial vein. However for interaction with a body, further modification of
the contact side is desirable by compounds favoring material attachment e.g. by RGD
proteins.
36
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8. LIST OF SHORTCUTS AND SYMBOLS
c crystallinity
AAL alanine-alanine-lysine
AFM atomic force microscopy
ATR-IR attenuated total reflectance infrared spectroscopy
BASYC® bacterial synthesized cellulose trademark
BD 1,4-butanediol
BC bacterial cellulose
BDI tetramethylene diisocyanate or 1,4-diisocyanatobutane
CT computed tomography
CNC cellulose nanocrystals
CNC-PEG cellulose nanocrystals grafted with poly(ethylene glycol)
CNF cellulose nanofibrils
DBTL dibutyltin dilaurate
DMA dynamical mechanical analysis
DMF N,N-Dimethylformamide
D-PHI degradable polar hydrophobic ionic polyurethane
DSC differential scanning calorimetry
DVO divinyl oligomer
B tensile strain at break or extensibility (%)
ELP4 elastin-like polypeptide-4
Ey Young's modulus (MPa)
FTIR Fourier transform infrared spectroscopy
GA glycolic acid
H12MDI dicyclohexylmethane diisocyanate
HA hydroxyapatite
HA hyaluronic acid
HDI hexamethylene diisocyanate or 1,6-diisocyanatehexane
LA lactic acid
LDI lysine diisocyanate
46
MDI Methylene diphenyl diisocyanate or 4,4'-methylene diphenyl diisocyanate
NCO/OH isocyanate ratio or isocyanate index
PCL polycaprolactone
PCN poly(hexamethylene carbonate)
PCU polycarbonate-urethane
PEG polyethylene glycol
PEO-PPO-PEO polyethylene oxide-co-polypropylene oxide-co-polyethylene oxide
PEU polyether-urethane
PGA polyglycolic acid
PHA polyhydroxyalkanoates
PHB polyhydroxybutyrate
PHV polyhydroxyvalerate
PLA polylactic acid
PLGA polylactic-co-glycolic acid
PP polypropylene
PU polyurethane
Q water uptake (%)
B tensile stress at break or toughness (MPa)
SBF simulated body fluids
SEM scanning electron microscopy
SMC smooth muscle cell
TE tissue engineering
TEDA triethylenediamine
TGA thermogravimetric analysis
U.S. FDA United States Food and Drug Administration
WAXS Wide-angle X-ray scattering
w/w weight/weight
wt.% weight percent
47
9. APPENDIX
Text is not available in the public version.
10. LIST OF FIGURES
Figure 1: General formation of polyurethane.
Figure 2: Modelled primary structure of segmented polyurethane [20].
Figure 3: Hydrogen bonding between the chains of a polyurethane based on 1,6 – hexane
diisocyanate and 1,4 – butane diol depicted according to [20].
Figure 4: Chemical structure of a) dibutyltin dilaurate and b) stannous octoate.
Figure 5: Reactions of isocyanate with other functional groups.
Figure 6: Schematic showing hypothesized mechanisms of PU scaffold degradation and
soluble breakdown products. (A) Hydrolysis of ester bonds to yield -hydroxy
acids (dotted line). (B) Abstraction of the -hydrogen atom in the R group
adjacent to the urethane bonds by reaction with hydroxyl radicals to yield lysine
or ethanolamine and carbon dioxide. (C) Abstraction of the -hydrogen atom in
lysine by reaction with hydroxyl radicals to yield multiple degradation products,
including NH4, -ketoacids, oximes, CO2, and carboxylic acids (degradation
products not shown to simplify the diagram) [63].
Figure 7: Schematics of (a) single cellulose chain repeat unit, showing the directionality of
the 1 -> 4 linkage and intrachain hydrogen bonding (dotted line), (b) idealized
cellulose microfibril showing one of the suggested configurations of the
crystalline and amorphous regions, and (c) cellulose nanocrystals after acid
hydrolysis dissolved the disordered regions. Figure depicted according to [97].
11. LIST OF TABLES
Table 1: Applications of polyurethanes in tissue engineering divided by U.S. FDA
standards [11].
Table 2: Isocyanates used in polyurethane chemistry for tissue engineering.
Table 3: Physical properties of the biodegradable polymers used as polyols in formation of
polyurethane.
48
Research activities
Articles
Vojtová, L., Kupka, V., Žídek, J., Wasserbauer, J., Sedláček, P., & Jančář, J. (2012).
Biodegradable polyhydroxybutyrate as a polyol for elastomeric polyurethanes. Chemical
Papers, 66(9), pp. 869-874.
Kupka, V., Vojtová L., Jančář, J., Solvent Free Synthesis and Structural Evaluation of
Biodegradable Polyurethane Films Based on Poly(ethylene glycol) and Poly(caprolactone),
Manuscript.
Competitions
2012/12 competition/conference: Student Competition and
Conference, Faculty of Chemistry, BUT, Brno. Awarded
presentation: Polyurethane Elastomers with Potential Use
in Medicine. The best presentation in doctoral section voted
by participants.
2013/6 competition: Ceitec Ph.D. competition, Brno, final round.
Presentation: Solvent Free Synthesized Polyurethane
Elastomers. Participation in the final round.
2014/5 competition: Jean-Marie Lehn’s Prize in Chemistry, Prague,
final round. Presentation: Polyurethane Elastomers with
Potential Use in Medicine. Participation in the final round.
Teaching
2012/autumn semester Practical class in structural analysis of polymers (master degree
students)
2014/15 Supervision of High school scientific activity. Student: Alžběta
Kryštofová from Řečkovice Grammar School, Brno. Topic:
Hydrolytical Stability of Biodegradable Polyurethanes