Upload
others
View
0
Download
0
Embed Size (px)
Citation preview
Acknowledged by the PhD committee and head of Hannover Medical School
President: Prof. Dr. Christopher Baum (until 31st December, 2019)
Prof. Dr. Michael Manns (as of 1st January, 2019)
Supervisor: Prof. Dr.-Ing Birgit Glasmacher (Institute for Multiphase
Processes, Leibniz Universität Hannover)
Internal Guide: Dr. Oleksandr Gryshkov (Institute for Multiphase Processes,
Leibniz Universität Hannover)
Co-Supervisors: Dr. Robert Zweigerdt (Leibniz Research Laboratories for
Biotechnology and Artificial Organs (LEBAO), Hannover
Medical School)
Prof. Dr. Sotirios Korossis (Mechanical, electrical and
manufacturing engineering, Loughborough University)
External expert: Prof. Eyal Zussman (Technion - Israel Institute of
Technology)
Internal expert: Prof. Dr. rer. nat. Andrea Hoffmann (AG Gradierte Implantate
und Regenerative Strategien, Medizinische Hochschule
Hannover Klinik für Orthopädie)
Day of public defence: 25th January, 2019
PhD project funded by:
The Federal Ministry of Education and Research as well as the Deutsche
Forschungsgemeinschaft (DFG, German Research Foundation) for the Cluster of
Excellence REBIRTH (EXC 62/3 valid until Dec 2017, EXC 62/4 valid until Oct 2019).
IP@Leibniz of Leibniz University Hannover promoted by the German Academic
Exchange Service (DAAD) for two exchange projects (one incoming, one outgoing)
during the study - project code 57156199.
ACKNOWLEDGEMENT
First off, I’d like to thank Prof. Dr.-Ing Birgit Glasmacher for giving me the
opportunity to do my thesis in the Institute for Multiphase Processes and guiding me
through this challenging time in my career. I would like to thank REBIRTH for
accepting me into their programme and funding my PhD through the DFG. I am also
grateful to Dr. Gerald Dräger and the rest of the IP@Leibniz team for funding two
exchange projects during the time of my PhD.
Sasha (Dr Oleksandr Gryshkov) has been an excellent mentor to me in academic
matters and also a great friendly comfort during stressful times. He constantly taught me
new things (sometimes without even meaning to), gave me honest criticism (which I
greatly appreciate) and gave me bear hugs and high fives when I did something well. I
am really lucky to have had him mentor me during my PhD. I would also like to thank
Dr Marc Müller for all our electrospinning brainstorming sessions, beer brewing
sessions and general PhD/career advice. It has been a pleasure working with him from
the very beginning.
I want to thank my co-supervisors Dr Korossis and Dr Zweigerdt (for their
expert scientific guidance), Rosi (for her administrative support), Almer (for his magical
computer fixing skills), Katja and Julia G (for their laboratory assistance), Prof. Wolkers
(for his assistance with the FTIR), the Institute for Technical Chemistry (for use of their
tensile testing machine) and Dr Pelz (for her constant support from REBIRTH).
I want to extend my gratitude to Dr Dagmar Pfeiffer, Dr Ingrid Lang and Julia
Fuchs and the rest of their department in Graz for accepting me into their laboratory and
providing me with support during my exchange project. I really did learn a lot during
that time. I’d also like to thank Ms Anne Marie Beck from NIFE for spending many
hours with me on histological stainings.
I would also like to thank my IMP colleagues who provided me with an
excellent work environment. Sara K has been the best office mate and a great friend. I
could always count on her to be there for me. I feel very fortunate to have had some
really motivated students: Mythili, Sarah K and Sarah C. It was a pleasure doing
research with them and building long lasting friendships in the process. I hope Sarah C
and I will have more ‘in the name of Science!’ moments in the future.
I do not have enough words to thank Irene and Jaspreet for being the best friends
I could have asked for. My time in Hannover was filled with laughter, positivity and
silliness because of them. I’d like to thank Andi for always being so kind to me. I cannot
express how grateful I am to have Terry in my life. He has made me the happiest person
in the world. Last but never least, I am ever thankful that I have such a supportive
family. My mother, father and sister are the reason I am where I am today.
ABSTRACT
Electrospun scaffolds are widely used in tissue engineering for the repair of soft
tissue. In spite of their obvious advantages, there are some practical limitations for in
vivo translation. One of the most crucial concerns is the lack of cell infiltration in such
scaffolds. To improve cell infiltration it would be ideal to simply increase pore size in
the scaffold. However, in electrospinning, this would also mean increasing fibre
diameter, which hampers cell functions such as attachment, proliferation and motility.
To circumvent this tradeoff, scaffolds with both small and large fibres have been
researched. In addition, the use of composite scaffolds (with synthetic and natural
polymers) have also been used extensively to improve scaffold-cell integration in vitro.
The research reported in this thesis, therefore, aims to improve upon existing
traditionally electrospun scaffolds by fabricating a hybrid scaffold which not only has a
large fibre diameter distribution, but is also a composite.
PCL-gelatin hybrid scaffolds were produced by altering the setup orientation of
electrospinning. The reasoning behind the change in microstructure in the vertical setup
orientation beyond a critical concentration of gelatin is presented in a hypothesis. The
characterisation and validation of these scaffolds were done through uniaxial
mechanical testing, a degradation study and the assessment of cellular infiltration (along
with a validation of cell viability and metabolic activity). The hybrid scaffolds were
mechanically stable and cell infiltration was improved in them while preserving cell
viability and metabolic activity. This thesis therefore offers a promising new method to
produce electrospun scaffolds with enhanced microstructure and cell infiltration for
application in soft tissue engineering.
TABLE OF CONTENTS
List of Figures 1
List of Tables 4
Abbreviations 5
Chapter 1 – Introduction
1.1 Tissue Engineering 7
1.2 3D engineering of soft tissue 10
1.3 Mimicking the Extracellular Matrix 12
1.4 Electrospinning - A method to fabricate fibrous scaffolds 13
1.5 Challenges in scaffold electrospinning 15
1.6 Cell interaction with micro/nanofibres 17
1.7 Enlarging pore size in electrospun scaffolds for enhanced cell infiltration 19
1.8 The fibre-pore paradox 19
1.9 Multimodal fibre diameter distributions in electrospun scaffolds 20
1.10 Composite scaffolds 22
1.11 The influence of setup orientation on electrospun scaffold microstructure 23
1.12 Rationale for the choice of polymers 25
1.13 Rationale for the choice of cells 27
1.14 Motivation and aim of the thesis 28
Chapter 2 – Materials and Methods
2.1 Electrospinning of PCL and PCL-gelatin scaffolds 31
2.1.1 Preparation of solutions 31
2.1.2 Solution properties 32
2.1.3 Electrospinning machine setup 32
2.1.4 Parametric optimisation of spinning 33
2.1.5 Imaging of electrospun scaffolds 35
2.1.6 Measurement of fibre diameter, pore size and thickness 35
2.2 Assessment of mechanical properties 36
2.2.1 Uniaxial tensile test setup 37
2.2.2 Representation of data 38
2.3 Degradation study 39
2.3.1 Experimental setup 39
2.3.2 Visualisation of fibrous structure 39
2.3.3 Raman analysis of gelatin loss and crystallinity of PCL 39
2.4 Cell response study 41
2.4.1 Sterilisation of electrospun mats for cell culture 41
2.4.2 Thawing and cultivation of cells 41
2.4.3 Cell seeding on scaffolds 42
2.4.4 SEM visualisation of seeded cells 44
2.4.5 Infiltration study 44
2.4.6 Viability study 45
2.4.7 Metabolic activity study 46
2.5 Statistical representation of data 48
Chapter 3 – Results
3.1 Electrospinning of PCL and PCL-gelatin scaffolds 49
3.1.1 Solution properties 49
3.1.2 Imaging of electrospun scaffolds 50
3.1.3 Measurement of fibre diameter and pore size 52
3.2 Assessment of mechanical properties 56
3.3 Degradation study 57
3.3.1 Visualisation of fibre structure 57
3.3.2 Raman analysis of gelatin loss and crystallinity of PCL 57
3.4 Cell response study 60
3.4.1 SEM visualisation of seeded cells 60
3.4.2 Infiltration study 61
3.4.3 Viability study 64
3.4.4 Metabolic activity study 66
Chapter 4 – Discussion
4.1 The influence of solution properties on electrospinning 67
4.2 The influence of setup orientation and polymer concentration on PCL-gelatin
blend electrospinning
70
4.3 Assessment of mechanical properties 77
4.4 Degradation study 79
4.5 Cell response study 80
Chapter 5 – Conclusion
5.1 Summary 85
5.2 Outlook 88
Appendix
References 91
Supplementary data 101
Publication list 108
Curriculum Vitae 109
Statement of contribution 111
Declaration 114
List of Figures
List of Figures
Fig 1.1 The scaffold-based tissue engineering approach has three main
components - a scaffold, cells and biological factors.
9
Fig 1.2 Cell-Cell and Cell-ECM adhesions are mediated by cell surface
proteins called Cell Adhesion Molecules (CAMs).
12
Fig 1.3 Formation of the Taylor cone during electrospinning. 15
Fig 1.4 Schematic depicting fibre-fibre contact points per unit length of
micro- and nanofibre.
17
Fig 1.5 Mechanism of cell-fibre binding in electrospun scaffolds. 18
Fig 1.6 Mesoscopically ordered scaffolds. 21
Fig 1.7 Gravitational and electric field directions in a) horizontal and b)
vertical setup orientations.
24
Fig 1.8 Taylor cone distortion. 25
Fig 1.9 Structure of a) PCL and b) gelatin. 26
Fig 1.10 Optimal fibre diameter and pore size requirements for proper
cell-scaffold integration.
29
Fig 2.1 Horizontal electrospinning setup. 32
Fig 2.2 Vertical electrospinning setup. 33
Fig 2.3 Schematic depicting how a) fibre diameter and b) pore size
measurements were made.
36
Fig 2.4 Size of the sample cut out for tensile testing 37
Fig 2.5 Uniaxial tensile test setup. 38
Fig 2.6 Scaffold cutting with custom made cutters. 42
Fig 2.7 Schematic depicting the steps in cell seeding. 43
1
List of Figures
Fig 3.1 SEM image panel showing microstructure of electrospun
scaffolds.
51
Fig 3.2 SEM image panel showing microstructure of electrospun
scaffolds made from the blends chosen for further analysis.
52
Fig 3.3 a) Fibre diameters and b) pore sizes measured from SEM images
of the electrospun scaffolds
53
Fig 3.4 Fibre diameter distribution 54
Fig 3.5 Pore size distribution 55
Fig 3.6 Mechanical properties. 56
Fig 3.7 Visualisation of fibre degradation over 30 days. 57
Fig 3.8 Gaussian fitted graphs of gelatin Amide I and PCL C=O. 58
Fig 3.9 SEM visualisation of cells seeded on the electrospun scaffolds. 60
Fig 3.10 Visualisation of infiltrated 3T3 cells on the electrospun scaffolds
over the span of 15 days.
61
Fig 3.11 Visualisation of infiltrated aMSCs on the electrospun scaffolds
over the span of 15 days.
62
Fig 3.12 Cell infiltration in the fibre mats over the span of 15 days. 63
Fig 3.13 Cell viability assay fluorescence images over the span of 7 days. 64
Fig 3.14 Cell viability percentages in the fibre mats over the span of 7
days.
65
Fig 3.15 MTT absorbances of the cells in the fibre mats over the span of 7
days
66
Fig 4.1 Taylor’s experiments on vertical jets of viscous fluids 71
Fig 4.2 Balancing forces involved during jetting. 72
Fig 4.3 The influence of the strength of the electric field on the start
point of jet instabilities.
73
2
List of Figures
Fig 4.4 Schematic representing the experimental results from Reneker
and Yarin.
74
Fig 4.5 Fibre orientation during tensile stretching. 78
Fig 4.6 Fibre interconnections and “nanowebs” observed in the vertically
spun blends.
79
Fig 5.1 Cryosection of a tubular electrospun PCL-gelatin scaffold seeded
with Human Iliac Artery Endothelial Cells (HIAECs) in the
lumen.
89
S.1 Thickness optimisation of electrospun scofflds 101
S.2 Mechanical testing Force-Elongation graphs (all sets) 102
S.3 Raman spectra - PCL pellets, gelatin powder 103
S.4 Degradation study - Raman spectra (fingerprint region) 104
S.5 FTIR spectra of PCL175V and PCL100g75V before and after UV
sterilisation
106
S.6 Scaffold cutter SolidWorks drawings 107
3
List of Tables
List of Tables
Table 1.1 Diversity in regenerative approaches for different target organs. 8
Table 2.1 Solution spinning parameters. 34
Table 2.2 Optimised seeding volumes and seeding densities for cell
response studies.
44
Table 3.1 Solution properties 49
Table 3.2 Summary of peak fitting results 59
Table 4.1 Some approaches to increase cell infiltration in electrospun
scaffolds
84
4
Abbreviations
Abbreviations
2D Two Dimensional
3D Three Dimensional
aMSC Amnion derived Multipotent Stromal Cell
AMSC Adult Multipotent Stem Cell
ANOVA Analysis of Variance
ATDV Aerodynamic Tangential Drag Vector
CAM Cell Adhesion Molecule
CI Confidence Interval
CL Collagen
cLSM Confocal Laser Scanning Microscope
CO2 Carbon dioxide
DMEM Dulbecco’s Modified Eagle Medium
DNA Deoxyribonucleic acid
ECM Extracellular Matrix
EDTA Ethylenediaminetetraacetic Acid
ESC Embyonic Stem Cell
EVF Electric Vector Field
FBS Foetal Bovine Serum
FDA Food and Drug Administration
FTIR Fourier Transform Infrared
FWHM Full Width at Half Maximum
GFV Gravitational Field Vector
H Horizontal
HA Hydroxyapatite
HCL Hydrochloric Acid
HIAEC Human Iliac Artery Endothelial Cells
5
Abbreviations
IMP Institute for Multiphase Processes
iPSC Induced Pluripotent Stem Cell
Mn Number average Molecular Weight
MTT 3-(4,5-Dimethylthiazol-2-yl)-2,5-Diphenyltetrazolium Bromide
Nd-YAG Neodymium-doped Yttrium Aluminum Garnet
NIFE Niedersächsischen Zentrum für Biomedizintechnik, Implantatforschung und Entwicklung
NP-40 Nonidet P-40
PBS Phosphate Buffered Saline
PCL Poly(ε-caprolactone)
PDGF-BB Platelet Derived Growth Factor with two B subunits
PEO Poly(ethylene oxide)
PLA Poly(lactic acid)
PLGA Poly(lactic-co-glycolic acid)
PLLA Poly(l-lactic acid)
PRP Platelet-rich Plasma
PVA Poly(vinyl alcohol)
PVC Polyvinyl chloride
PVDF Polyvinylidene Fluoride
RGD Arginine-Glycine-Aspartate
SEM Scanning Electron Microscope
SF Silk Fibroin
Si Silicon
TE Tissue Engineering
TFE 2,2,2 - Trifluoroethanol
UV Ultraviolet
V Vertical
6
Chapter 1: Introduction
maintain their mechanical integrity until the remodelling process is complete (15).
Furthermore, there is a great stress on the porosity, pore interconnectivity and surface to
volume ratio of the scaffold for efficient cell colonisation (21,22).
Based on the application, a plethora of fabrication techniques have been
employed to create porous, biodegradable and mechanically stable scaffolds that can
guide neotissue formation. Some of these techniques include solvent casting, particulate
leaching, gas foaming, self assembly, phase separation, electrospinning, rapid
prototyping, melt moulding and freeze drying (23). Each of these techniques have their
own merits and demerits and it is important to select the scaffold fabrication method
depending on the properties demanded by the end application.
1.2 3D ENGINEERING OF SOFT TISSUE
The term ‘soft tissue’ refers to any tissue that is not hardened or calcified. In
particular, it is tissue that surrounds bone or internal organs and may play a connective
or supportive role (24). This category encompasses tissues such as hair, cartilage, nerve,
muscle, skin, fat, fascia etc. Soft tissue damage is incredibly common in daily life,
caused either by genetic defects, trauma, disease or ageing. The gold standard for
treating this sort of damage is currently the use of autologous implants. However, the
main pitfall is that autologously implanted tissue is easily absorbed and rapid losses in
volume result in only about 40-60% viable cells. In addition, donor site morbidity is a
pervasive problem in autografting preventing its widespread use (25). Given the
inherent low regenerative abilities of soft tissue, tissue engineering has emerged as a
feasible option for such restoration.
The biological mechanisms that drive cellular function have been studied
predominantly through two-dimensional (2D) tissue culture. However, this does not
reflect the intricately structured three-dimensional (3D) labyrinth of the extracellular
matrix (ECM) in which cells exist in vivo. The topography of the cellular
microenvironment (the immediate environment perceived by the cells) directly
influences the phenotype of cells and can modify behaviours such as migration,
10
Chapter 1: Introduction
differentiation and mechanotransduction. This has urged researchers to simulate the in
vivo 3D environment in scaffolds for better functionality of cell-seeded constructs (26).
3D porous scaffolds not only provide a significantly more biomimetic template for cell
growth, but also enable enhanced biological signalling (cell-cell as well as cell-scaffold)
and better mass transport (27).
In light of this, a great deal of research has been done in 3D reconstruction of
damaged soft tissue such as skeletal muscle (28), skin (29), nerves (30), blood vessels
(31), cornea (32), ligament (33), trachea (34), adipose tissue (35) and heart valves (36),
etc. A wide variety of synthetic and natural polymers have been employed to
accomplish this, for example - poly(ε-caprolactone) (PCL), poly(vinyl alcohol) (PVA),
poly(lactic acid) (PLA), chitosan, collagen, alginate, gelatin, etc (37).
Although the basic concept of tissue engineering is straightforward, the practical
realisation is remarkably complex. As such, there are some limitations to restoring
volume loss in soft tissue. Organ patterning is one of the primary challenges thus far.
The scaffold has to allow for structural and functional mimicry of native tissue and also
have a construction conducive for vascular ingrowth. A high surface to volume ratio is
difficult to achieve, especially taking into consideration oxygenation and mass transport
of nutrients and waste, but is important for long-term cell survival. Microenvironment
considerations are of tremendous importance, involving the control of spatiotemporal
distribution of biological factors, facilitating adherence and infiltration of neotissue.
Dynamic matrix modification is required to carefully guide cell migration, maximising
the infiltration capacity of the scaffold. It is also important to match the mechanical
properties of the scaffold to the anatomical implantation site, particularly in load
bearing applications. Inadequate mechanical strength can lead to implantation failure in
the long run, reducing the ability of the tissue engineered scaffold to bond with native
tissue. Furthermore, there are limitations on the depths up to which regeneration can
still occur. Several factors may play a role in this, such as resistance to cell infiltration,
unsuitable mechanical properties and inadequate vascular supply. It is imperative to
keep these criteria in mind while designing a soft tissue replacement for successful
implementation of the same in vivo. (38)
11
Chapter 1: Introduction
feedback mechanism where the ECM influences cellular response (‘outside-in
signalling’), which in turn can alter the ECM itself (‘inside-out signalling’). Fabricating
such multi-faceted ECM analogues to replicate its structure, function and physiological
cell-ECM interactions is therefore a daunting task. (39) Ideally, an artificial scaffold
would execute all the roles of native ECM and mimic its complexity and diversity.
While this seems impossible today, current advances in scaffold design show that it is
possible to replicate at least some of these functions through the addition of nanofibres
and the incorporation of natural polymers (39).
Nanofibres in the range of ECM fibre diameters can dramatically improve cell
function because the microenvironment scale closely resembles that of the ECM. While
synthetic polymers have been successfully used to replicate this nanofibrous structure
and mechanical support, they lack specific cell interaction domains to promote cell
adhesion and migration. On the other hand, natural polymers enhance bioactivity and
biocompatibility but are difficult to process and lack mechanical stability. (27) Recent
efforts to combine synthetic and natural polymers to form composite fibrous scaffolds
have shown better results with respect to scaffold properties and cell-scaffold
integration (42-46).
1.4 ELECTROSPINNING - A METHOD TO FABRICATE FIBROUS
SCAFFOLDS
Currently, electrospinning, self-assembly and phase separation are the three
techniques used to synthesise nanofibres. Of these, electrospun scaffolds are the most
widely studied and seem to be immensely promising for application in tissue
replacements (47,48-52). The basis for the discovery and conceptualisation of
electrospinning dates back to the sixteenth century when William Gilbert first described
the phenomenon of electrostatic attraction (53). Since then, research on the distortion of
fluid droplets under the influence of electrostatic forces has been studied in great detail.
Anton Formhals made significant patent contributions to the field of electrospinning
between 1931 and 1944 (54-61). Following this, Sir Geoffrey Ingram Taylor
mathematically modelled the conical droplet shape (now referred to as the Taylor cone)
13
Chapter 1: Introduction
formed by fluids under the influence of an electric field in 1969 (62). Since the early
1980s, the concept of electrospinning has been widely researched in both textile as well
as tissue engineering (63). In an attempt to mimic the complex 3D network of polymer
and polysaccharide fibres (50-500 nm) of the native extracellular matrix (ECM) of
adherent cells, electrospinning has now emerged as the preferred fabrication method for
nano and microfibrous scaffolds.
Electrostatic fibre spinning, or ‘electrospinning’, involves the stretching of a
viscoelastic solution into a fibre of sub micron diameter that is collected in a random
manner on a grounded collector. In general, a polymer is solubilised in a suitable
conductive solvent at a certain concentration and extruded through a needle. A collector
is placed some distance from the needle and is grounded. When the nozzle is connected
to a high voltage a strong electrostatic potential difference is developed between the
nozzle and the collector, causing charges to collect on the surface of the polymer
solution droplet. An increase in voltage at the nozzle causes an increase in charge
density and repulsion, resulting in droplet distortion (Taylor cone)(Fig 1.3). At a critical
voltage, the molecules overcome the surface tension of the droplet and the polymer
solution is drawn into a single fibre in either the nano or microrange. At some distance
from the nozzle, the jet becomes unstable and a whipping motion causes the fibre to
stretch and thin and the solvent rapidly evaporates. By the time the fibre reaches the
collector, it is deposited dry (under optimised parameters) in a random manner, forming
a dense fibrous mesh. (64)
Electrospinning presents significant advantages over other types of fibre
fabrication methods because of its versatility and tunability. Fibre morphology and pore
size can be controlled by modifying the many parameters involved. Depending on the
intended replacement, collector shape and size can also be modified. Furthermore, it is
possible to use a wide variety of materials ranging from purely biological to synthetic to
ECM analogous materials. This flexibility in material usage allows for targeted
mechanical strength, degradation rate and bioactive components (16). Electrospun
scaffolds also have a large void volume, interconnected porous network and high
surface to volume ratio conducive for the incorporation of cellular components (65).
14
Chapter 1: Introduction
effect of electrospinning. It is unclear if this really is a cause for concern. While there
are some reports detailing the detrimental effects of solvent retention (67), there are
others that claim that the amount of residual solvent is too low to lend toxicity in vitro
or in vivo (68). The choice of material has a strong influence on how much solvent is
retained in the scaffold. Even in the case of some retention, they can be removed with a
simple heat and vacuum treatment (69). Nevertheless, this aspect is important to keep in
mind. Another constraint in electrospinning is the limit to the thickness of electrospun
scaffolds. This is because of two reasons. First, long processing times cause the build-up
of charges and prevent further deposition of fibres (70). Second, the insulation of the
collector increases with an increase in thickness of the deposited material. Hence, there
is a difficulty in repairing large tissue defects.
The chief pressing concern in electrospun scaffolds is insufficient cell
infiltration. Pore sizes in nanofibrous scaffolds are much smaller than general cell
diameters and therefore do not allow cells to infiltrate easily. Optimum pore sizes,
although not standardisable for all cell types, have been suggested to be in the range
100-500 µm. Electrospun scaffold pore size is far below this requirement, resulting in
limited vascular ingrowth, hypoxia and inefficient mass transfer. Additionally, the
packing density of fibres in electrospun scaffolds is high because each fibre layer is
continuously deposited over the previous layer while still being strongly attracted to the
collector. This results in some degree of compression, thus reducing the void volume of
the scaffold (66).
Electrospun scaffolds show a direct relationship between fibre diameter and pore
size, which means that smaller fibre diameters increase the incidence of fibre-fibre
contact points which in turn reduces pore area (Fig 1.4). Notwithstanding the high
porosity, the inadequate pore size restricts cellular penetration. Therefore, despite the
potential of these scaffolds to provide a 3D microenvironment comparable to that of
native ECM, most electrospun constructs in reality behave as 2D surfaces (66,71,72).
As highlighted previously, it is prudent to cultivate cells in a cell-permeable 3D
construct because it leads to better cell-cell signalling as well as more space for cells to
respond to mechanical cues. Poor cell infiltration can result in an inhomogeneous
16
Chapter 1: Introduction
1.7 ENLARGING PORE SIZE IN ELECTROSPUN SCAFFOLDS FOR
ENHANCED CELL INFILTRATION
The electrospinning community has invested much effort into enlarging the pore
size of scaffolds and attempting to solving the problem of insufficient cellular
penetration. The simplest solution attempted by many research groups is the
straightforward manipulation of parameters such as solution flow rate, applied voltage,
tip to collector distance and solution properties (81-83). An increased flow rate,
decreased voltage, shorter tip to collector distance and higher solution viscosity all
result in thicker fibres and consequently larger pores. However, it is a well known fact
that cells perform better in nanofibrous microenvironments than in their microfibrous
counterparts (as detailed in section 1.6). This has led to a much higher incidence of
electrospun nanofibres with other modifications for increasing pore size. Using
sacrificial material is an extremely popular strategy to increase pore size. Scaffolds with
two polymers, one immediately dissolvable and the other a slow degrading component,
have been fabricated. The fast degrading component is dissolved away or leached out
after electrospinning to create void spaces in the scaffold. The sacrificial element can be
a polymer solution that has been co-spun with the main polymer (84-87), salts (salt
leaching) (88) or ice crystals (cryogenic electrospinning) (89). Gas foaming is also a
well researched technique (90) and is sometimes combined with salt leaching (91).
Other methods that have been employed to increase pore size include melt
electrospinning (92), wet electrospinning (93), modification of collector shape (94,95),
combination of electrospinning with other fabrication technologies (such as direct
polymer melt deposition) (96) and laser ablation (97). Simultaneous polymer
electrospinning and electrospraying of cells has been employed to improve cell density
in electrospun scaffolds without changing the pore size (98).
1.8 THE FIBRE-PORE PARADOX
Keeping in mind overall cell health and success of tissue integration, there are
two opposing practical requirements for cells cultivated on electrospun scaffolds, which
brings us to the crux of the matter. Fibre diameter and pore size of electrospun scaffolds
19
Chapter 1: Introduction
are closely related, directly interdependent entities. Pore size increases with increase in
fibre diameter and vice versa. It is important to have a large pore size for adequate cell
infiltration (which invariably increases the fibre diameter), but it is also important to
have nanofibres that enhance cell adhesion and proliferation (which decreases pore
size). Owing to the interdependence of pore size and fibre diameter, it is difficult to
fabricate a scaffold with nanofibres and macropores (99,100). Pore size modifications
allow for larger pores but do not solve the problem of inefficient cell adhesion or loss in
mechanical stability. For instance, the use of sacrificial elements may increase pore size,
but the sudden increase in void volume can drastically hamper the mechanical
properties of the scaffold. Certain soft tissue replacements in load bearing sites such as
muscle and cartilage require high biomechanical strength. Electrospun scaffolds
currently are already unable to provide the optimum properties for high physical stress
regions. Loss of bulk material only serves to worsen the problem. Nanofibrous scaffolds
on the other hand have a high surface to volume ratio (100) but have little to no cell
infiltration and better function as patterned membranes for cell adhesion.
1.9 MULTIMODAL FIBRE DIAMETER DISTRIBUTIONS IN
ELECTROSPUN SCAFFOLDS
In order to overcome the fibre-pore paradox in electrospinning, there have been
massive efforts to combine nano and microfibres in the same scaffold. A number of
studies have attempted a trade-off situation where both micro and nanofibres are
present. Some noteworthy examples are elucidated in this section.
Kidoaki et al. proposed two methods for mesoscopically ordered scaffolds -
multilayering electrospinning and mixing electrospinning (Fig 1.6). Multilayering
electrospinning was achieved by sequentially spinning layers of different polymers.
They also proposed a vascular graft model with the multilayering concept. Moving a
step further from the simple juxtaposition of different fibre diameters, mixing
electrospinning was done by spinning two separate polymer solutions, with optimised
parameters for two different fibre diameter outputs, simultaneously on the same
collector (101).
20
Chapter 1: Introduction
1.10 COMPOSITE SCAFFOLDS
The concept of hybrid scaffolds in terms of scaffold composition is not a new
one. It has been proven time and time again that combining structural and bioactive
components in one scaffold afford it superior functionality to those made of only one
material. There are numerous ways to achieve this, but the most researched strategy is
the combination of synthetic and natural polymers. Since this research is focused on
electrospun scaffolds, this review will explore the functionality of blend electrospun
scaffolds, coated scaffolds and scaffolds incorporated with biological agents.
Synthetic biodegradable polymers have been extensively employed in tissue
engineering applications because of the ability to tailor their mechanical properties,
control the degradation rate and easily shape them into various configurations. They are
also favoured because of the gamut of polymers available for processing, compatible
with a number of fabrication techniques. Out of these, polyesters are particularly
attractive because of their hydrolysis based degradation, producing degradation
products that in most cases are eliminated safely through metabolic pathways. It is also
possible to tailor degradation rates to the application by altering the structure (105).
However, they are largely hydrophobic and lack ligands for successful cell binding
(106). Synthetic polymers are also sometimes preferred because they are cheap and can
be obtained in large quantities with negligible batch to batch variation (47).
Natural polymers on the other hand, are polymers derived from natural sources
and can be classified as proteins, polysaccharides and polynucleotides. The major
advantage of these polymers is their hydrophilicity and ability to provide an ECM-like
microenvironment by presenting ligands for cell binding. Biological factors and natural
polymers are able to induce some degree of cell infiltration by a phenomenon called
‘chemotaxis’ where cells respond and migrate toward chemical signals (39,66). This can
be done by introducing chemical gradients, soluble signals or even electrical potentials
(66). As such, the problem of poor cell infiltration can be addressed in part without
changing the scaffold microstructure. Unfortunately, natural polymers in general lack
mechanical stability and can vary in composition from source to source making it
difficult to maintain repeatability in experiments (47).
22
Chapter 1: Introduction
To capitalise on the advantages and compensate for the disadvantages of both
types of polymers, composite scaffolds have been developed and tested. The easiest way
to incorporate natural polymers into synthetic scaffolds is by surface coatings.
Numerous papers have expanded on the positive influence of surface functionalisation.
Coatings such as heparin (107), gelatin (108), collagen (109), fibrin (110), decorin (111)
and chitosan (112) have been successfully applied. Blend electrospinning of synthetic
and natural polymers has allowed for preservation of mechanical properties while
lending biocompatibility and hydrophilicity to the scaffolds. Examples include PCL-
collagen (113), PCL-gelatin (114,115), PLGA-gelatin-elastin (116), etc. Synthetic
polymers have also been incorporated with biological factors like platelet-rich plasma
(PRP) (117) and chemotaxis agents such as platelet-derived growth factor (PDGF-BB)
(118).
1.11 THE INFLUENCE OF SETUP ORIENTATION ON ELECTROSPUN
SCAFFOLD MICROSTRUCTURE
Electrospinning apparatuses are usually either setup horizontally or vertically (in
this case, top-down systems are preferred over bottom up systems due to ease of setup) .
In horizontal electrospinning, the nozzle is situated parallel to the ground and the
collector is placed perpendicular to this. The electric vector field (EVF) is always
between the nozzle and collector, so in this case it is parallel to the ground. The
gravitational field vector (GFV) obviously never changes and acts perpendicular to the
ground. In vertical electrospinning the EVF and GFV are both in the same direction
(perpendicular to the ground) because the collector is placed parallel and the nozzle is
placed perpendicular to the ground (115) (Fig 1.7). Either setup can be used as the
process is similar, but horizontally arranged systems are sometimes preferred because
the occurrence of artefacts in the final product is minimised (119).
The influence of setup orientation on the microstructure of the final electrospun
product has hardly been investigated. Yang et al. (120) electrospun PVDF in three
orientations - horizontal, shaft (top-down vertical) and converse (bottom-up vertical).
They observed that the fibre diameters obtained for the same range of voltages was
23
Chapter 1: Introduction
response. (134) Gelatin is hydrophilic, but it is also a polyelectrolyte. It contains amine
and carboxylic groups that are easily ionised in aqueous solution. Combined with its
strong hydrogen bonding, it becomes a very difficult material to electrospin. Gelatin is
also very quickly degraded without cross linking and tends to lose mechanical integrity
under physiological conditions. It is therefore common for gelatin to be blended with
stable synthetic polymers when fabricating scaffolds for tissue engineering. (135)
1.13 RATIONALE FOR THE CHOICE OF CELLS
Tissue engineering has seen the use of differentiated primary cells, cell lines and
stem cells. Differentiated primary cells, although advantageous in many ways, have
disadvantages such as the invasive nature of tissue extraction, limited expansion
potential ex vivo and the development of senescence (136). In place of this, cell lines are
commonly used as they are cost effective, easy to handle, can be indefinitely expanded
in culture and do not have the ethical concerns that primary cells are associated with.
Cell lines are consistent and the experimental results are reproducible but they usually
do not fully replicate the behaviour of primary cells because they are genetically
modified and may present altered phenotypes and stimuli responses. (137) Stem cells on
the other hand, are of three types - embryonic stem cells (ESCs), adult multipotent stem
cells (AMSCs) and induced pluripotent stem cells (iPSCs). Embryonic stem cells face
ethical issues and iPSCs raise concerns due to modified genotypes which makes
AMSCs currently key candidates for regenerative therapies (138). Two cell types were
considered for the validation of the electrospun scaffolds in this thesis - NIH 3T3 cell
line and amnion derived Multipotent Stromal Cells (aMSCs).
The NIH 3T3 mouse embryonic fibroblast cell line was first isolated in 1962 by
George Todaro and Howard Green. Since then, they have been used in numerous studies
because they are robust and have a predictable growth pattern (139,140). 3T3 cells are
quite commonly used in proof of concept studies since the results are repeatable and
comparable. However, the use of cell lines cannot be used to simulate the behaviour of
‘normal’ cells in the body. The use of 3T3 cells especially has been hotly debated due to
its metastatic tendencies under certain conditions (141). It is for this reason that while it
27
Chapter 1: Introduction
is acceptable to perform preliminary cell response studies with 3T3 cells to deem a
scaffold safe for cell colonisation, further validation must be done with sensitive
primary cells.
The MSCs used in this project were derived from the placental amnionic
membrane of the common marmoset (Callithrix jacchus) by the Institute for
Transfusion Medicine, Hannover Medical School (PD Dr. Thomas Müller) and stored at
-150 °C prior to use. These cells are excellent candidates for preclinical primate (non-
human) studies because of their non-invasive extraction, easy availability, plasticity and
immunosuppression abilities. Amnion MSCs have been shown to have higher
proliferation capacity in comparison to their bone marrow counterparts making them
preferable for expansion (138,142). MSCs are also sensitive cells that respond strongly
to surface topography, scaffold composition and mechanical stress, making them an
interesting choice for comparison with 3T3 cell function.
1.14 MOTIVATION AND AIM OF THE THESIS
Many different cell types (usually application oriented) have been used to
validate different electrospun scaffolds, with varying degrees of success. However, it is
interesting to note that there is a tacit understanding among members of the
electrospinning community that electrospun scaffolds have to be tailored to the cell
type. This is understandable as each cell type has its own distinct ECM and for
optimised outcomes, the scaffold should possess components particular to that native
site. While this statement definitely rings true, inadequate cell infiltration is a common
problem in all electrospun scaffolds. In particular, enhancing cell infiltration while also
preserving cell viability and metabolism is a universal challenge. It is therefore
necessary to provide a common solution.
In view of the general state of the art described in the previous sections, it is
clear that it is crucial to have many opposing scaffold elements for different
requirements. The first and foremost is the necessity to have large interconnected pores
for sufficient cell infiltration (provided by large microfibres). The second is the need to
28
Chapter 2: Materials and Methods
Table 2.1 - Solution spinning parameters (115). PCL and PCL-gelatin blend solutions were
electrospun in the horizontal and vertical setup orientations using optimised parameters.
The four final categories of scaffolds were chosen for characterisation due to
three main reasons. First, PCL175V was chosen as the unblended category to observe
the effect of gelatin addition. Second, PCL125g50 is the only blend that produced
scaffolds with homogenous nanofibres (spun in the horizontal orientation) and a
gradient of fibre diameters from nano to micro range (spun in the vertical orientation.
Both these categories were chosen to see the effect of changing fibre diameter and pore
size, keeping the blend concentration the same. And last but not least, PCL100g75V
was chosen because it has a similar scaffold architecture to PCL125g50V but has a
higher concentration of gelatin, giving us an idea of how gelatin concentration can
affect the properties of the scaffold and cell infiltration.
It was initially debated if the varied fibre structure obtained in the PCL-gelatin
scaffolds was caused due to electrospinning in two different instruments (unfortunately,
the fibre mats could not be spun horizontally and vertically in the same machine due to
the large size of the collector). However, this effect was verified by electrospinning
tubes (on a much smaller collector) in both setup orientations in the same device where
similar variations in fibre structure were observed in PCL-gelatin scaffolds above a
critical concentration of gelatin (data not shown).
Name of
the blend
Polymer
concentration
(mg/ml)
Mass ratio of
PCL:gelatin
Horizontal (H) and Vertical (V)
electrospinning parameters
PCL GelatinVoltage
(kV)
Flow rate
(ml/h)
Working
distance
(cm)
H V H V H V
PCL175 175 0 1:0 23 20 3 3 27 26
PCL150g25 150 25 6:1 13 15 2 2 19 22
PCL125g50 125 50 5:2 13 13 1.5 2 19 22
PCL100g75 100 75 4:3 x 13 x 2 x 22
PCL100g100 100 100 1:1 x x x x x x
34
Chapter 2: Materials and Methods
2.1.5 Imaging of electrospun scaffolds
Prior to imaging, dry scaffolds were sampled into small pieces and mounted
onto sample holders with conductive double-sided adhesive tape. They were then
sputter coated in vacuum with palladium using the EMITECH SC7620 sputter coater
(Quorum Technologies). Plasma sputtering times for samples of 50 µm thickness was
30 s and for samples of thickness 150 µm it was 45 s. The samples were then imaged
using the Hitachi S-3400 N with EDAX Scanning Electron Microscope (SEM) system
at a magnification of 1000x (for the purpose of image paneling), an accelerating voltage
15 kV and working distance of 7 mm. Samples were also imaged for the purpose of
measuring fibre diameter and pore size at different magnifications according to the scale
of the sample. For thickness measurements only, samples were cut in liquid nitrogen to
get a clear cut edge without compressive damage. Samples were then mounted on
vertical sample holders with the cut edge facing upwards and imaged under the SEM.
2.1.6 Measurement of fibre diameter, pore size and thickness
All morphological measurements were made using Fiji Image J software. For
fibre diameter and pore size, triplicate samples were imaged per blend for statistical
significance. Five images were taken per sample, five fibres/pores were chosen at
random per image and five measurements were made per fibre/pore, yielding a total of
375 measurements per blend (115) (Fig 2.3). For thickness estimation, 10 measurements
were taken per sample and then averaged.
Pore size measurement of electrospun scaffolds has been fiercely contended in
the scientific community. The problem arises because electrospun scaffolds do not have
true pores, there are gaps that are simply created by the continuous layering of fibre
meshes. Since a 3D pore is difficult to categorise, most research groups resort to
measuring the 2D inter-fibre distances, created by fibres in the top couple of layers,
visible on an SEM image. Although it is possible to visualise many layers beneath the
surface, measurements on the resultant 2D images are only accurate for surface fibres.
The ‘pore sizes’ graphed in the corresponding results section are also 2D inter-fibre
distances measured from SEM images of the same samples (thickness 50 µm) used for
fibre diameter measurement.
35
Chapter 2: Materials and Methods
2.3 DEGRADATION STUDY
In order to assess the stability of the samples, fibre degradation and gelatin loss
induced by cell culture medium in vitro, a degradation study was performed. This was
accomplished in two ways. Firstly, SEM was done to visualise any morphological
changes in the fibres during the course of the study. Secondly, Raman spectroscopy was
performed to analyse the chemical composition of the scaffolds during the length of the
degradation study. Undegraded and degraded samples (for 15 and 30 days) were
analysed alike. All samples were of 150 µm thickness.
2.3.1 Experimental setup
Sample preparation for SEM and Raman spectroscopy was performed in a
similar manner. All four scaffold samples were placed in DMEM (pH 7.4) in a
humidified incubator at 37 °C and 5 % CO2 for a span of 30 days. Prior to imaging,
samples were first washed in bi-distilled water and then dehydrated in an increasing
gradient of ethanol concentrations (30 %, 50 %, 70 %, 100 %) for 10 min each. This
was followed by air drying.
2.3.2 Visualisation of fibre structure
Triplicate samples for SEM imaging were sputter coated and imaged at a
magnification of 1000x, an accelerating voltage 15 kV and working distance of 10 mm.
Here as well, five images were taken per sample. The fibre structure was then visually
examined for defects such as thinning or breakage.
2.3.3 Raman analysis of gelatin loss and crystallinity of PCL
The dried constructs were placed on a glass slide using adhesive tape and
imaged using a Raman Microscope (Alpha 300 RA, WITec GmbH). Prior to
measurement, a Si wafer was used to calibrate the system as the Si peak is highly stable
at 520 cm-1 (corresponding to the crystalline Si-Si bond longitudinal optical phonon
vibrations (149)). 532 nm excitation light from an Nd-YAG Laser was focused on the
scaffold samples at a magnification of 100x. Acquisition of Raman spectra was done in
39
Chapter 2: Materials and Methods
5 s at a laser power of 15 mW. Spectra were acquired from 5 different fibres of varying
thicknesses (above 3 µm).
Following this, analysis of data was carried out using Project FIVE.plus (WITec,
Ulm). The final spectrum was stitched to show only the fingerprint region of the data
(800-1800 cm-1). The results were graphed and standard background subtraction was
performed (there were no artefact peaks of cosmic ray interference). The intensity
values were normalised using formula 2.2:
Iy - Imin
Inorm = —————— (2.2)
Imax - Imin
where,
Inorm - normalised intensity
Imax - maximum intensity value
Imin - minimum intensity value
Iy - measured intensity
After normalisation, the data obtained was correlated with known spectra and
characteristic peaks of PCL and gelatin were identified (refer supplementary S.3 and S.4
for all Raman spectra). The Amide I peak of gelatin was deconvoluted using Origin,
integrated for area under the curve and then percentage loss of gelatin (from one tested
day to another) was calculated using equation 2.3. PCL crystallinity was analysed by
measuring the ratio between the full width at half maximum (FWHM) of amorphous
and crystalline peaks.
A2 - A1
L = —————— x 100 % (2.3)
A1
where,
L - percentage loss
A1 - area under curve 1
A2 - area under curve 2
40
Chapter 2: Materials and Methods
2.4 CELL RESPONSE STUDY
2.4.1 Sterilisation of electrospun mats for cell culture
Scaffolds were sterilised by UV-C radiation (254 nm) (150) using the UV lamp
in the cell culture work bench (151). Fibre mats were laid out on aluminium foil and
exposed to UV radiation for 30 min (15 min each side). They were then rolled up and
stored in sterile 50 ml Falcon tubes and used for cell seeding under aseptic conditions
thereafter. Fourier Transform Infrared spectroscopy (Perkin Elmer Spectrum 100 FTIR
spectrometer) was performed on unsterilised and sterilised samples to ensure that UV
radiation did not alter or denature the chemical composition of the scaffolds. The
wavenumber range was chosen as 4000 to 650 cm-1 and the number of scans was set to
8. The probe force value was 50 for all analysed samples (see supplementary S.5 for
spectra and associated analysis).
2.4.2 Thawing and cultivation of cells
3T3 cells were cultivated in growth medium prepared by mixing 500 ml of
Dulbecco’s modified Eagle’s medium (DMEM) (Biochrom GmbH) that contains 3.7 g/
L of NaHCO3, 4.5 g/L of D- Glucose, stable glutamine, Na-Pyruvate and low endotoxin
with 75 ml of foetal bovine serum (FBS) (Biochrom GmbH) and 5.8 ml of Penicillin/
Streptomycin (Biochrom GmbH). aMSCs were cultivated in the same medium with the
addition of 580 µL of ascorbic acid. 3T3 cells were thawed at Passage 3 and aMSCs
were thawed at Passage 7 for subsequent propagation. For all experiments, aMSCs were
seeded at passage 9 to ensure comparable proliferative capacity.
Cryopreserved cells were thawed by swirling the frozen vial swiftly in a warm
water bath at 37 °C. The vial was then placed in an icebox and further handled under
aseptic conditions. The cell suspension was then transferred to a falcon tube (Sarstedt)
and 5 ml of cold medium at 4 °C was added dropwise to the tube. This was then
centrifuged at 4 °C for 5 min at a speed of 1000 rpm. Following centrifugation, the
supernatant was aspirated and the cell pellet was resuspended in 1 ml growth medium.
The entire contents were then plated on a 60.1 cm2 coated culture dish (TPP AG) for
further passaging.
41
Chapter 2: Materials and Methods
3T3 cells were passaged at 80% and aMSCs were passaged at 70 % confluence.
All solutions used during passaging were warmed to 37 °C before use. The medium was
aspirated and cells were washed with 1xPhosphate Buffered Saline (PBS) (pH 7.4,
Biochrom GmbH) and cells were detached by the addition of 0.05 %/0.02 % (w/v)
trypsin-EDTA (Biochrom GmbH) and incubation for 3 min at 37 °C. Trypsin activity
was stopped by adding medium and cells were resuspended over the dish surface
several times and collected. After centrifugation for 5 min at 1000 rpm, the cell pellet
was isolated and further resuspended in medium. Cell membrane integrity was assessed
by the Trypan Blue (Sigma Aldrich) exclusion method using the Vi-CELL XR Cell
Viability Analyser (Beckman Coulter). Pre seeding viabilities were always > 90 %.
Cultivation was continued in a humidified incubator at 37 °C with 5 % CO2 and
medium was changed every couple of days. Cells were then passaged on confluency.
2.4.3 Cell seeding on scaffolds
All electrospun constructs were seeded in uncoated 12-well plates with an area
of 5 cm2 (Sarstedt) to allow the cells to preferentially attach to the scaffolds over the
plate surface. Circular cutouts were made with custom cutters (manufactured at the IMP
workshop, see supplementary S.6 for drawing) with diameters of either 1.6 cm or 0.6
cm depending on the assay (Fig 2.6).
Fig 2.6 - Scaffold cutting with custom made cutters. Two diameters were used in the cell
response study depending on the assay - 1.6 cm and 0.6 cm.
All scaffold samples were wetted from the bottom with a little medium. This
serves two important purposes - first, it prevents the samples from being sucked into the
airflow mechanism of the cell culture bench, and second, it pre-wets the scaffolds to
allow for quick cell attachment. Cells were aliquoted into the required seeding
42
Chapter 2: Materials and Methods
Table 2.2 - Optimised seeding volumes and seeding densities for cell response studies.
2.4.4 SEM visualisation of seeded cells
All chemicals mentioned in this section were purchased from Carl Roth GmbH.
The morphology of adherent cells on the scaffolds and their interaction with the fibres
were visualised using SEM imaging. Cells were seeded on scaffolds using the
previously discussed cell seeding procedure (seeding days were considered as day 0).
On days 1 and 5, prior to fixation, the samples were first rinsed with 0.1 M Cacodylate
buffer (prepared from Cacodylic acid sodium salt trihydrate at pH 7.4). Fixation was
performed by incubating cells for 15 mins with 2.5 % Gluteraldehyde in 0.1 M
Cacodylate buffer (pH 7.4). After another washing step, they were placed in bi-distilled
water for 15 min and subsequently dehydrated in an ascending gradient of Ethanol (30
%, 50 %, 70 %, 99 %). This was followed by air drying, sputter coating and SEM
imaging as described in 2.1.5.
2.4.5 Infiltration study
The cells were allowed to grow and proliferate on the scaffolds for a span of 15
days. Infiltration depths were assessed for all the four blends on days 1, 5, 10 and 15.
Cells were then stained with Phalloidin (λex 495 nm, λem 520 nm) (152) and Hoechst
33342 (λex 346 nm, λem 460 nm) (153) and visualised using a confocal laser scanning
microscope (cLSM) from Carl Zeiss GmbH (LSM 510). Both stains were purchased
from Sigma Aldrich. Phalloidin stains F-actin filaments green and Hoechst stains the
cell nuclei blue.
Scaffold
diameter
Seeding
volume
Scaffold
thickness
Cell seeding
densityCharacterization assay
1.6 cm 100 µl 150 µm ~ 50,000 cells/cm2
Live/dead assay
Infiltration depth measurement
SEM visualisation of seeded cells
0.6 cm 25 µl 150 µm ~ 25,000 cells/cm2 MTT assay
44
Chapter 2: Materials and Methods
For staining, fixation of cells on the scaffold was first performed by rinsing the
cells twice with 1xPBS and subsequent incubation with 4 % formaldehyde (Sigma
Aldrich) for 15 min followed by washing with PBS. After fixation, cell membrane
permeabilisation was done by incubation for 15 min in 0.5 % Triton-X (Sigma Aldrich).
Cells were then washed with PBS, stained with Hoechst (1:1000 dilution factor) and
Phalloidin (2:1000 dilution factor) for 45 min in the dark and washed again with PBS to
remove any excessive stain. Samples were mounted on glass sides using Mount
Fluorcare (Carl Roth GmbH) mounting medium at room temp and left to dry overnight.
Samples were imaged inverted in the cLSM. Laser lines Argon/2 for the 458-514
nm range and Diode 405-30 for the 405 nm channel along with the Plan- Neofluar 20x/
0.5 objective (with 0.17 mm cover glass correction) were used to image the scaffolds.
Pinhole was set to 110 corresponding to an optical slice thickness of 5 µm. Green and
blue excitation channels were used to view the actin filaments and to locate nuclei
respectively. Focus was gradually fine-tuned and the gain was adjusted to avoid over
exposure while imaging. Images were acquired in Z-stacks (of thickness 5 µm). Total Z-
stack depth was noted for each sample and the resulting images were saved as .lsm files.
2.4.6 Viability study
Cell viability was assessed using a two-color assay (Calcein AM and Ethidium
homodimer-1 both purchased from Sigma Aldrich) to distinguish between live and dead
cells. Non-fluorescent cell permeant Calcein AM is converted to fluorescent calcein (λex
496 nm, λem 516 nm) in living cells indicating intracellular esterase activity. Ethidium
homodimer-1 is normally cell membrane impermeable and enters dead cells indicating
loss of plasma membrane integrity and gives a red (λex 528 nm, λem 617 nm) fluorescent
signal on binding to nucleic acids (154).
Viability study was performed on day 1, 4 and 7 of culture. Working solutions of
Calcein AM (1 mg/ml, 1:1000) and Ethidium homodimer – 1 (1 mM, 2:1000) were used
to prepare staining solutions in 1xPBS for scaffolds. The staining solution was
thoroughly vortexed before use.
45
Chapter 2: Materials and Methods
Pre seeded scaffolds were retrieved from their respective cell culture plates and
were washed twice with 1xPBS solution. Then, 500 µl of staining solution was added to
each well and the scaffolds were incubated in the dark for 20 min. Finally, they were
washed twice with 1xPBS to remove any excessive stain and viewed inverted under the
Axiovert 200M fluorescence microscope (interfaced with Zeiss Axiocam MRm camera,
power source MAC 500, X-cite Series 120 fluorescence excitation lamp and computer
with Axiovision imaging software). Five images were taken at 10x magnification per
sample (middle, top, right, bottom and left).
For cell quantification, the image analysis software Image J (Fiji) was used.
Analysis was performed by splitting the blue and green channels. The ‘find maxima’
function was used to count the number of live and dead cells. Noise tolerance was
adjusted in increments of 5 until the background staining was excluded. Calculations
were also verified manually to account for errors.
The total cell number was determined using formula 2.3 and the percentage of
live cells was calculated using formula 2.4:
Nt = Nlc + Ndc (2.4)
Nlc
V = ——— x 100 % (2.5)
Nt
where,
Nt - total cell number
Nlc - number of live cells
Ndc - number of dead cells
V - viability percentage
2.4.7 Metabolic activity study
The MTT assay is a colorimetric method to assess the metabolism of viable
cells. Viable metabolising cells in culture reduce MTT (3-(4,5-dimethylthiazol-2-
yl)-2,5-diphenyltetrazolium bromide) to purple formazan which possesses an
46
Chapter 2: Materials and Methods
absorbance maximum near 570 nm. Since the metabolic activity (and by extension the
rate of MTT reduction) of cells can depend on a myriad of factors (without exactly
affecting viability), it is not, as often incorrectly interpreted, a cell proliferation
indicator (155). For this study, MTT was purchased from Sigma Aldrich.
Cells were seeded on scaffolds as described in section 2.4.3. Controls (scaffolds
without cells) were run separately for each blend of scaffold for all the measured days.
Cells were cultured on the scaffolds for 7 days and readings were taken on day 1, 4 and
7 in order to assess the proliferation rate over time. Scaffolds were transferred to a 96
well plate (TPP AG) for the assay. The solubilising solution (0.04 M HCL (Sigma
Aldrich), 0.1 % NP-40 (Applichem GmbH) in isopropanol (Carl Roth GmbH)) and
MTT stock solution (5 mg/mL) in PBS were prepared prior to the experiment.
On each test day, the cell culture medium was first aspirated, followed by the
addition of 10 µl of MTT solution and 90 µl of serum free media to each well. Serum
free media was used in order to avoid any interference during absorption readings. The
plate was then incubated in a humidified incubator at 37 °C with 5 % CO2 for 4 hours.
The MTT solution was then aspirated and 100 µL of solubilising solution was added to
each well followed by incubation at 37 °C with 5 % CO2 for 1 hour. After 1 hour, 50 µL
of the solution was transferred from each well to another 96 well plate for measuring
absorbance values using a multi-plate reader at 570 nm. 50 µl of solution was taken for
absorbance readings to ensure uniformity in testing, primarily to account for solvent
absorption by the porous scaffold.
For calculations, control values were subtracted from the actual values obtained
from the scaffolds to obtain corrected sample absorbances (which were graphed) using
formula 2.5:
Absc = Abssam - Abscon (2.6)
where,
Absc - corrected absorbance
Abssam - absorbance of the sample
Abscon - absorbance of the control
47
Chapter 3: Results
3.1.2 Imaging of electrospun scaffolds
After the initial optimisation of the electrospinning parameters (Table 2.1), all
scaffolds were prepared at a thickness of 50 µm for visualisation of scaffold
microstructure under the SEM. SEM images are panelled in Fig 3.1 with the
horizontally spun samples on the left side and the vertically spun samples on the right
side.
Fibre diameters were measured (n=125) for PCL 175 (Fig 3.1a and Fig 3.1b) and
it was observed that the vertically spun sample (1.52 ± 0.08 µm) showed a marginally
higher variation than in the horizontally spun sample (1.66 ± 0.07 µm). Pore size
measurements also did not differ statistically between the horizontal (20.3 ± 1 µm) and
vertical (23.9 ± 1.6 µm) orientations. Fig 3.1c and 3.1d represent the microstructure of
the PCL150g25 blend in the horizontally and vertically spun orientations respectively.
Here as well, there is no observable morphological difference between the two
scaffolds. Both setup orientations produce a similarly homogeneously structured
scaffold.
However, when the gelatin concentration is increased, i.e in PCL125g50 (Fig
3.1e and 3.1f), we see a stark difference between the horizontal and vertically spun
samples. The sample spun in the horizontal setup is homogeneous and consists mostly
of nanofibres but the sample spun in the vertical setup is heterogeneous and consists of
a range of fibre diameters (even though the blend concentration is similar in both cases).
Note the large degree of variation in pore size within PCL125g50V itself. Pore size is
21.4 ± 1.4 µm in PCL125g50H versus 115 ± 0.08 µm in PCL125g50V. The effect is
sustained when the gelatin concentration is increased further in PCL100g75V (pore size
107 ± 9 µm) but produced a very unstable electrospinning in the horizontal setup.
Therefore, we see a variation in scaffold microstructure in the vertical orientation above
a critical concentration of gelatin in scaffolds spun from blends with the same overall
concentration of solutes (175 mg/ml).
An attempt was made to spin PCL100g100 but the process was unstable in both
orientations yielding no proper samples for further testing.
50
Chapter 3: Results
Fig 3.1 - SEM image panel showing microstructure of electrospun scaffolds. Thickness of
scaffolds = 50 µm. All images were taken at 1000x magnification. Scale bar = 50 µm. Image
has been reused modified from (115).
51
Chapter 3: Results
Mean fibre diameters are 1.52 ± 0.05 µm (PCL175), 0.59 ± 0.04 µm
(PCL125g50H), 5.15 ± 0.42 µm (PCL125g50V) and 4.52 ± 0.33 µm (PCL100g75V).
Fibre diameter distributions are represented as histograms graphed in Fig 3.4.
Note that the range of fibre diameters is much larger in the vertically spun samples,
especially in PCL125g50V where the bin centres run from 0.5 to 17.5 µm. The range of
fibre diameters obtained here is much larger than attainable in a normal electrospinning
process.
Fig 3.3 - a) Fibre diameters and b) pore sizes measured from SEM images of the electrospun
scaffolds. n=375, **** p<0.0001, whiskers represent min and max values, mean shown as ‘+’.
Image 3.3a has been reused modified from (115).
Pore sizes of all four samples are depicted in Fig 3.3b. The results correlate with
the fibre diameter measurements. The mean pore size of PCL125g50V (115 ± 5 µm) is
almost four and a half times that of PCL175V (25.1 ± 1.1 µm) and four times that of
PCL125g50H (27.5 ± 1.3 µm). Similarly, the mean pore size of PCL100g75 is also high
(111 ± 5 µm). Pore size distributions are graphed in Fig 3.5.
53
Chapter 3: Results
Fig 3.4 - Fibre diameter distribution. Bin centres for PCL175V and PCL125g50H have been
truncated at 3.5 µm as there are no measurements beyond that size for those samples. Bin sizes
for all are 1 µm. n=375.
54
Chapter 3: Results
Fig 3.5 - Pore size distribution. Bin centres for PCL175V and PCL125g50H have been
truncated at 90 µm as there are no measurements beyond that size for those samples. Bin sizes
for all are 20 µm. n=375.
55
Chapter 3: Results
3.2 ASSESSMENT OF MECHANICAL PROPERTIES
Mean force at break and mean strain at break are graphed in Fig 3.6.
As depicted in the graphs, we can see that the introduction of gelatin reduces the
strain of the samples making them more brittle. Mean strain percentage at break for the
gelatin blends are 39.2 ± 11.7 % (PCL125g50H), 46.4 ± 85.6 % (PCL125g50V) and
34.2 ± 61.7 % (PCL100g75V) against 180 ± 604 % for PCL175V.
However, the mean force at break was much higher in the vertically spun blends
with the hybrid morphology compared to the other two samples. Mean values of
PCL125g50V, PCL100g75V and PCL175V are 3.4 ± 0.08 N, 2.19 ± 0.64 N and 1.55 ±
2.34 N respectively. PCL125g50H despite having the same concentration as its
vertically spun counterpart, only shows a mean force at break of 1.29 ± 0.13 N.
Fig 3.6 - Mechanical properties. a) mean force at break, b) mean strain at break. Cross-
sectional area of samples = 1.5x0.005 cm, n=3.
56
Chapter 3: Results
3.3 DEGRADATION STUDY
3.3.1 Visualisation of fibre structure
Undegraded samples and samples degraded in cell culture medium (at 37 °C and
5 % CO2) for 15 and 30 days were visualised under the SEM (Fig 3.7). All samples
were of thickness 150 µm. Points of fibre degradation (obviously thinned regions on a
fibre) were observed only in the PCL175V sample. Vertically spun blends visibly
showed uniformly thinner fibres by day 30.
Fig 3.7 - Visualisation of fibre degradation over 30 days. Obvious points of fibre thinning
have been indicated by yellow circles. All images were taken at x1000 magnification. Scale bar
= 50 µm.
3.3.2 Raman analysis of gelatin loss and crystallinity of PCL
Raman spectra of all four samples were taken and analysed for gelatin loss and
crystallinity of PCL gelatin. Fig 3.8 shows the characteristic peak of gelatin Amide I at
1636-1668 cm-1 (157) and C=O stretch of PCL at 1730 cm-1. The C=O stretch in PCL is
actually composed of two constituent peaks - a narrow peak at 1725 cm-1 representing
the crystalline phase and a broader peak at ~1735 cm-1 representing the amorphous
phase (158).
57
Chapter 3: Results
Fig 3.8 - Gaussian fitted graphs of gelatin Amide I and PCL C=O. a) PCL175V, b)
PCL125g50H, c) PCL125g50V, d) PCL100g75V. Day 0 represents undegraded samples while
day 15 and day 30 represent degraded samples.
Areas under the Amide I peak of gelatin show a 36.99 ± 6.65 % loss in
PCL125g50H and a 53.44 ± 6.48 % loss in PCL100g75V by day 15 (Table 3.2). After
that, the gelatin loss seems to stagnate in both samples from day 15 to day 30.
PCL125g50V shows no significant loss in gelatin. This peak also shows a minor shift
toward higher wavenumbers from day 0 to day 30 in all the tested samples. Crystallinity
of PCL could not be calculated for samples using information from the entire spectrum
due to interference from the gelatin peaks. Further, FWHM ratios of the crystalline to
amorphous C=O peak showed no discernible trend in the tested period.
58
Chapter 3: Results
3.4 CELL RESPONSE STUDY
3.4.1 SEM visualisation of seeded cells
Seeded cells were fixed and visualised on day 1 and day 5 of culture. The
behaviour of 3T3 cells and aMSCs was comparable. On day 1, it can be seen that cells
reside on the surface of PCL175V and PCL125g50H owing to their low pore size.
However cells are seen between the first few layers of fibres in the vertically spun
blends. By day 5, this effect is well pronounced and some infiltration can be observed in
the vertically spun blends. The formation of monolayers on the other two samples show
that the cells are restricted from infiltrating the bulk of the scaffold. (Fig 3.9) Cells
seeded on the gelatin blends exhibited a much better spread of cytoplasm compared to
PCL175. The total coverage of the scaffold surface was also more. Specifically in
PCL125g50V, additionally to the infiltration, it can be seen that the cells extend their
philopodia to attach securely to the nanofibres in 3D.
Fig 3.9 - SEM visualisation of cells seeded on the electrospun scaffolds. a) 3T3 cells, b)
aMSCs. All images were taken at x1000 magnification. Scale bar = 50 µm. Image 3.9a has been
reused modified from (115).
60
Chapter 3: Results
3.4.2 Infiltration study
Cells cultured on the electrospun constructs were stained with Phalloidin and
Hoechst for imaging under the cLSM. Infiltration depths were measured as the z-stack
depth down to which cell nuclei are visible. Images showing the stained cells are
panelled in Fig 3.10 (3T3 cells) and Fig 3.11 (aMSCs). Since gelatin is strongly
autofluorescent, there is sometimes a background fluorescence depending on the
amount of gelatin leaching out.
Fig 3.10 - Visualisation of infiltrated 3T3 cells on the electrospun scaffolds over the span of
15 days. Scale bar = 50 µm.
61
Chapter 3: Results
Fig 3.11 - Visualisation of infiltrated aMSCs on the electrospun scaffolds over the span of 15
days. Scale bar = 50 µm.
In correlation with this qualitative analysis, the measured infiltration depths are
graphed in Fig 3.12. PCL175V and PCL125g50H show minimal infiltration in the first
few days and then a stagnation from day 5 onwards. On day 15, 3T3 cells cultured on
PCL175V showed a mean infiltration depth of 22.33 ± 6.7 µm, which is a negligible
increase from 18.3 ± 3.1 µm measured on day 1. aMSCs fair better in this case where
their infiltration depth increases from 17.3 ± 1.5 µm on day 1 to 37.3 ± 3.6 µm on day
15 on PCL175V. PCL125g50H shows the lowest infiltration of both 3T3 cells (22 ± 5.3
µm) and aMSCs (26.7 ± 2.2 µm) on day 15.
62
Chapter 3: Results
3.4.3 Viability study
The infiltration study shows that the cells have infiltrated (or not) through the
bulk of the scaffold but gives no indication of the viability of these cells. A live/dead
assay (Calcein AM for live cells and Ethidium homodimer-1 for dead cells) was
performed for this purpose on seeded cells on days 1, 4 and 7 (Fig 3.13).
Fig 3.13 - Cell viability assay fluorescence images over the span of 7 days. a) 3T3 cells, b)
aMSCs. Scale bar = 100 µm. Images PCL125g50V day 7 and PCL100g75V day 7 in 3.13a have
been taken from (115).
64
Chapter 4: Discussion
influence of strong electric fields and this can cause a broad divergence in fibre
diameters (159). Polyelectrolytes tend to greatly increase the conductivity of a solution
because they dissociate in aqueous media to produce charged ions. While this is a good
option to increase the conductivity of certain solutions to facilitate electrospinning
(160), an excess of these ions can hinder the formation of fibres.
Since gelatin is a polyelectrolyte and has gels in aqueous media, it cannot be
electrospun in water and is most often mixed with synthetic polymers and dissolved in
an organic solvent. TFE in particular has been suggested as a good solvent for
polypeptides (84). While the use of an organic solvent is helpful to make a
polyelectrolyte solution spinnable, the conductivity is still expected to be somewhat
high.
The gelatin blends used in this study show a much higher conductivity than the
pure PCL solution, but are electrospinnable. However, it was observed that an increase
in the concentration of gelatin resulted in more instabilities during electrospinning. As
such, it was not possible to obtain sufficiently good samples with a solution
concentration of PCL100g100, as the solution constantly collected and dried at the
needle tip and formed multiple jets during the process.
The electrospinning of a single polymer is predictable as the solution properties
do not vary tremendously. But there is a critical concern in blend solutions about phase
separations. More often than not, constituent polymers of blend solutions are not fully
miscible and require some dopant to facilitate this.
In this project, 2 wt% acetic acid is added with respect to TFE. Literature shows
that just 0.2 wt% acetic acid is sufficient to mediate proper gelatin dissolution in TFE
and miscibility with the PCL solution (148). However, this was not observed in the
context of this project in the long term. A 0.2 wt% acetic acid was found to allow for
better dissolution and initial miscibility but does not prevent phase separation after a
few hours. This could be problematic for electrospinning because some solutions need
to be spun for a few hours to get the required fibre mat thickness. Therefore, the acetic
acid dopant concentration was increased to 2 wt% and thereafter no phase separation
was observed even overnight.
69
Chapter 4: Discussion
The possibility of phase separation during electrospinning was also considered.
Since it is not possible to ascertain the incidence of phase separation during
electrospinning, a Raman analysis was carried out on the finished products to see if the
composition of the thick and thin fibres varied significantly (as part of the degradation
study, graphs in supplementary). The variation was found to be negligible with both
PCL and gelatin distributed more or less homogeneously through the fibres.
4.2 THE INFLUENCE OF SETUP ORIENTATION AND POLYMER
CONCENTRATION ON PCL-GELATIN BLEND ELECTROSPINNING
Electrospinning PCL175 in both orientations produces scaffolds with marginal
fibre diameter differences. A similar outcome was also reported by Yang et al. (120)
(albeit without statistical analysis) but the result reported here does not differ
statistically. Therefore we can say that PCL175 spun in either orientation has a similar
structure. Solution conductivity and viscosity are constant and a stable Taylor cone is
observed during the electrospinning process.
When gelatin is introduced at a low concentration, it does not affect the
electrospinning process and again, homogeneously structured scaffolds are obtained in
both setup orientations (PCL150g25). However, beyond this concentration of gelatin, a
gradient of fibre diameters is observed only in the vertical setup orientation (as seen in
PCL125g50V and PCL100g75V).
In blend electrospinning, variable fibre diameters have been associated with
improper solution mixing and phase separation of constituent polymers before
electrospinning (148). In the context of this project, no phase separation was observed
in any of the solutions before electrospinning and therefore there must be another reason
for the variable fibre diameters observed in vertically spun PCL-gelatin blends (above a
critical concentration of gelatin). While an extensively detailed analysis involving other
blends (with and without polyelectrolyte components) and complicated fibre
characterisation methods is beyond the scope of this study, it is possible to propose a
preliminary hypothesis to explain such a situation based on existing literature.
70
Chapter 4: Discussion
gelatin) of the working solution, there is a field-enhanced dissociation of the
contaminants which additionally causes the conductivity to rise exponentially
depending on the applied electric field.
Horning and Hendricks also examined the forces driving an electrostatic jet in
1979 (164), of which conductivity was one studied parameter. While they were unable
to explain how exactly the conductivity affects the geometry of the jet, they showed that
it indeed affected where on the jet the instabilities start to occur. As explained
previously this would eventually affect the resultant fibre diameter. In 2009, Stanger et
al (165) characterised the Taylor cone in terms of charge density. They claimed that an
increase in solution conductivity causes an increase in charge density at the Taylor cone
resulting in a smaller jet diameter (because the jet is drawn from a smaller effective area
or “virtual orifice”) and lower rate of fibre mass deposition.
This brings us to the topic at hand. There are three aspects to keep in mind in the
context of this thesis. The first crucial point to be noted is that Taylor’s experiments
were performed with conducting and non-conducting fluids, not with polymers
dissolved in solvent. Polymer molecules not only interact with each other, but also with
the needle tip and the electrostatic force provided by ions (162). The number of polymer
molecules flowing into the droplet as well as the effect of the interactions of these
molecules affects the local conductivity in the droplet. By extension, this will also affect
the charge density, jet diameter and resultant fibre diameter.
Second, we cannot predict the number of gelatin and PCL molecules in the
Taylor cone at any particular instant because of the general randomness of fluid flow
into the droplet. This may result in a small fluctuation in polymer ratio and consequent
fluctuation in the charge density at the Taylor cone. Now, if a non-polyelectrolyte was
used in place of gelatin, the fluctuation would be negligible. But due to the presence of
gelatin, the effect of the fluctuation is amplified because of its inherent strong ionisation
property. When the gelatin concentration is very low (as in the case of PCL150g25), it
will have no discernible effect on the final product. When the gelatin concentration is
increased, it may cause some variation in fibre diameters but may still be too weak on
its own to cause the huge range of fibre diameters obtained in this work.
75
Chapter 4: Discussion
Third, since the mass of polymer carried away by the jet reduces with increasing
charge density, the fluctuations in local conductivity at the Taylor cone could induce
fluctuations in the ‘ideal’ voltage for that system. This means that the range in which
there can be oscillations in droplet size and jet diameter is increased. The generally high
conductivity of gelatin prevents PCL-gelatin blends from being spun at high voltages
due to the formation of multiple jets. Therefore, the overall spinnable range is narrowed
while the oscillation range is widened.
This review finally brings us to the first part of the hypothesis that was
generated to help explain the phenomenon of varying fibre diameters in PCL-gelatin
scaffolds above a critical concentration of gelatin. It is proposed that the noticeable fibre
diameter variations are caused by a combination of the three aspects mentioned above.
But, we still encounter two unanswered questions. Is the combination of the above
aspects sufficient to cause such a large variation in fibre diameter as was observed? And
why does it only happen in the vertical setup orientation?
Since the only parameter that changed in both setup orientations is the direction
of the gravitational force, we can attempt to answer these questions with the last piece
of the puzzle - Taylor’s take on the correlation between the electric field and the start of
the jet instability. As elucidated before, the point of jet instability is pushed closer to the
needle when the electric field is stronger and is pushed away from the needle when the
electric field is weaker.
In the vertical orientation, the fluctuation in fibre diameter caused due to
changes in the charge density at the Taylor cone and the voltage being in the oscillation
range is exacerbated by gravity. A noticeable fluctuation in the jet instability starting
point was observed during electrospinning. The point moved up and down over a
certain range on the extruded jet. Since the fluid has nowhere to go but down to the
collector, the jet drew out more or less solution, became thick or thin and got deposited
on the collector depending on the condition at a given time.
In the horizontal system however, gravity is not directly involved in balancing
the aerodynamic drag and electric field. This means that the start point of the jet
instability cannot move much further away from the needle as there is no gravity to
76
Chapter 4: Discussion
support the fluid flow in that direction. A lowering of the electric field will just result in
the jet being unable to traverse the full distance to the collector. Additionally, it is
possible that the sagging of the Taylor cone (also reported in (121) and (122)) acts as a
buffer, preventing the charge density fluctuations from having any real impact on the jet
thickness or the resulting fibre diameter. Correspondingly, a barely noticeable shift back
and forth of the start of the jet instability was observed during electrospinning. This
would also well explain why PCL125g50H has a marginally higher fibre diameter
variation than PCL175 (leading to an inconsequential increase in pore size) while
PCL125g50V and PCL100g75V have a much larger fibre diameter variation (leading to
a substantial increase in pore size).
The instability in electrospinning PCL-gelatin solutions resulting in non-
uniformity and inconsistencies in the fibres with increasing gelatin content was also
reported by Nelson et al (166) although it was just a trivial secondary observation to
their main results.
Unfortunately, it was not possible to verify this hypothesis, given the time frame,
the goals and scope of the thesis. But perhaps an in depth analysis can be done in the
future on the electrospinning of blend solutions in both setup orientations with and
without a polyelectrolyte.
4.3 MECHANICAL PROPERTIES
When comparing the mechanical properties of scaffolds to human tissue, it is
important to keep in mind how the test was performed in both situations. There are
inconsistencies in literature about how uniaxial tensile tests are setup. They can be
either quasistatic or dynamic, and in either case, may have a different strain rate. Since
the failure strain is dependent on the type of test and the strain rate used in the study, it
is difficult to pinpoint a precise value for either a biological tissue or a scaffold. For
instance, abdominal porcine skin was reported to have a failure strain of 123-126 %
when using a quasistatic strain rate of 0.25-10 %/s in one paper (167) while it was
measured as 31-53 % when using a quasistatic strain rate of 50 mm/min in a another
77
Chapter 4: Discussion
reduced by the addition of gelatin. Since gelatin is brittle by nature, it was expected that
the strain percentage would be reduced. In addition, the mixing of PCL and gelatin is
only realised on a physical level (i.e. they are not copolymers). The physical bonds
between both polymers are much weaker than the covalent bonds in homopolymers and
hence, blending two different polymers on a physical level often tends to reduce the
mechanical performance (strain) (156).
The addition of gelatin greatly increased the mean force at break in the vertically
spun blends, especially PCL125g50V. The vertically spun blends present a certain
degree of interconnections in the very large fibres caused due to delayed evaporation of
solvent (Fig 4.6) which tend to increase the force at break (170-172). They also show
nanofibres bridging the microfibres creating “nanowebs” (Fig 4.6) as Soliman et al
(104) observed in their experiments.
Fig 4.6 - Fibre interconnections and “nanowebs” observed in the vertically spun blends.
4.4 DEGRADATION STUDY
The degradation study shows that PCL-gelatin blending is able to retain gelatin
in the scaffold without the need for crosslinking. After an initial drop in gelatin
concentration in the first 15 days (only in PCL125g50H and PCL100g75V), the
scaffolds tend to retain what is left through to day 30 without losing more.
79
Chapter 4: Discussion
PCL100g75V loses about half its gelatin content within the first 15 days only in cell
culture medium. Cell cultivation would likely increase the gelatin loss even more within
this time. Despite being of similar fibre structure to PCL125g50V, PCL100g75V poses
an obstruction to the infiltration of cells, as seen in the infiltration study. A similar initial
drop and then stabilisation of gelatin content in PCL-gelatin scaffolds was also observed
by Hwang et al (91).
In general, semi-crystalline PCL at 20 °C shows sharp peaks corresponding to
regions of alkyl chain segment or the ester group. These sharp peaks are an indicator of
the crystallinity of the sample. Broad low intensity peaks are often associated with
amorphous states indicating that the polymer has lost some of its conformational order.
(173) Typically, aliphatic degradable polymers such as PCL tend to show an increase in
crystallinity (and an associated larger stiffness) in the beginning stages of degradation
because of the preferential surface erosion of amorphous regions (127,174). However,
such an obvious increase in crystallinity due to the degradation of amorphous regions is
only expected after about four weeks of degradation (174). No trend of change in
crystallinity was discernible from the obtained results for the studied time period.
4.5 CELL RESPONSE STUDY
Traditionally, fibre diameter variations in electrospun products have been
considered undesirable and researchers strived to achieve uniform fibres in electrospun
scaffolds (148, 175). However, more recently, research into the functions of the ECM
has driven scientists to create more heterogeneous scaffolds in terms of structure as well
as composition. Cell response is a complicated phenomenon and is controlled by an
array of factors. The main aim of this part of the study was to show that the electrospun
scaffolds prepared were able to increase cell infiltration compared to traditionally
electrospun scaffolds over a given period of time, regardless of the cell type used. The
secondary aim of the cell studies was to show that the infiltrated cells show a good
viability and good metabolic activity.
80
Chapter 4: Discussion
Initial cell response studies were performed with 3T3 cells because the cells are
robust and their size, function and population doubling time (20-26 hours (140)) are not
largely affected from passage to passage like that of primary cells. 3T3 cells showed a
much higher infiltration in the vertically spun blends compared to PCL175V and
PCL125g50H where stagnation was observed. The cells cultivated on PCL175V and
PCL125g50H tended to form monolayers at the top of the scaffolds because they were
unable to penetrate through the dense fibrous structure with small pores. PCL125g50V
especially outperformed all the other scaffolds because of its high pore size.
The question arises here as to why PCL100g75V (having a similar structure to
PCL125g50V) did not perform just as well. However, this can be explained by the fact
that gelatin absorbs copious amounts of water in aqueous medium, swells and tends to
form a gel before it gradually leaches out from the scaffold. Since PCL100g75V is
almost 50 % gelatin, the excessive swelling and gelation can cause a physical barrier to
infiltrating cells. PCL125g50V on the other hand has enough gelatin to mediate
hydrophilicity and biocompatibility, but not too much that it hinders cell penetration.
3T3 cells showed an extremely good viability on all scaffold samples with an
increase in total cell number throughout the study period. Cell survival on day 1 was
comparable for all scaffolds. The average total number of cells were much higher in all
the gelatin samples compared to PCL175 on day 7 showing a higher proliferation of
cells in the gelatin blends. However, the viability was specifically found to be higher by
day 7 only in PCL125g50V and PCL100g75V. Since PCL125g50H has a similar
composition to PCL125g50V but does not show the same high viability (despite having
a comparable total cell number on day 7), the better performance of PCL125g50V can
be attributed to its microstructure. The larger pores of the vertically spun blends likely
allowed for better transport of oxygen, nutrients and waste, leading to better cell
viability.
The assessment of metabolic activity gives us an indication of how healthy the
viable cells are. It is not only important that the cells have a high viability percentage, it
is also crucial that they metabolise at a sufficient rate. The MTT assay can show
increased absorbances if either the viable cell number is high (but each cell on its own
81
Chapter 4: Discussion
may not metabolise to its full extent), the viable cells are metabolising at a very high
rate (without a noticeable increase in cell number), or the cells are large in number and
metabolise at a high rate. Since the 3T3 cells have a very small population doubling
time, it is difficult to deduce whether the increased metabolic activity over the days is
only because of the increased number of cells or because the cells are also increasing
their metabolism.
MSCs on the other hand behaved very differently to the 3T3 cell line. It is well
known that MSCs are very sensitive to microstructure, scaffold composition and
mechanical properties. Unlike the 3T3 cells, they do not grow indiscriminately on any
suitable surface and have a population doubling time much higher than that of 3T3 cells
(176), especially after passage 6 (177). Population doubling time of MSCs from the 6th
to 12th passage varies between ~60 and ~160 h depending on the passage number,
source species (human or non-human) and the site of extraction (amnion, bone marrow,
umbilical cord, placenta or adipose tissue) (177, 178, 179).
The infiltration study showed a similar result to that observed with the 3T3 cells.
Cells cultivated on PCL125g50V showed the highest mean infiltration depth by day 15.
PCL100g75V, here as well, was not able to compete with PCL125g50V presumably
because of excessive gelatin swelling. PCL125g50H was the only sample that showed a
stagnation in infiltration depth from day 10. Cell viability percentages in the fibre mats
were quite different to what was seen with the 3T3 cells. Cells cultivated on PCL175V
showed a significant reduction in viability from day 1 to day 4 that did not recover by
day 7. The reduction in viability in PCL175V is caused because PCL is hydrophobic
and does not have a large pore size for infiltration and mass transport, presenting an
overall unfavourable growth environment for aMSCs. While all the samples with
gelatin show no significant difference between each other on all three days, it is
interesting to note that PCL125g50H has a significantly lower viability on day 7
compared to its viability on day 4. PCL125g50H, in contrast to PCL175V, does contain
gelatin but also does not have large pores. This effect of suboptimal growth conditions
could not be seen in the 3T3 cells because they are very tough. aMSCs therefore
provide a much more realistic simulation of how cells would behave in vivo.
82
Chapter 4: Discussion
Since the aMSCs were seeded on the scaffolds at passage 9, it was not expected that
they proliferate at a high extent. Apart from cells seeded on PCL175V (where the
viability significantly dwindled), the total number of cells observed during the viability
study from day 1 to day 7 in all the gelatin blends was more or less constant. Therefore
the result obtained in the MTT assay (either an increase or decrease in absorbance) is a
function of cell metabolic rate and not change in population.
PCL125g50H showed the highest metabolic activity on all the tested days
because it contains gelatin and also has an excellent nanofibrous surface. PCL125g50V
and PCL100g75V show a lower metabolic activity on day 1 (because of the addition of
large microfibres) but quickly catch up to PCL125g50H by day 7. Now this is intriguing
because while the cells initially are exposed to both nano and microfibres (on the
surface), they are able to navigate through the large pores and seek out nanofibres for
attachment. PCL125g50V is the only sample that shows an increasing trend in cell
metabolic activity. Metabolic activity is linked to cells that have a higher motility and
therefore it is understandable that as the cells infiltrate more, they also increase their
metabolic activity. The generally improved behaviour of the aMSCs on the
heterogeneously structured gelatin blended scaffolds can also be a result of increased
scaffold stiffness as it is known that stem cells are highly mechanosensitive (as
evidenced by improved attachment, motility and differentiation capabilities) (180,181).
From the results obtained in the cell study in terms of infiltration, viability and
metabolic activity, it is clear that a heterogeneous fibre morphology is an asset in
electrospun scaffolds facilitating infiltration as well as preserving cell viability and
metabolic activity. The addition of gelatin largely increases the biocompatibility of the
scaffolds and greatly influences cell function. However, it is important to regulate the
ratio of gelatin to PCL so that the swelling of gelatin does not hinder the movement of
cells. Therefore, we propose that PCL125g50V has tremendous potential for further
studies and application in soft tissue engineering. An overview of some other attempts
to increase cell infiltration in electrospun scaffolds is shown in Table 4.1.
83
Chapter 5: Conclusion
These scaffolds were also characterised in terms of mechanical properties and
stability during degradation. To briefly recapitulate, the addition of gelatin seemed to
reduce the strain percentage of the scaffolds. However, the incidence of
interconnections between fibres together with the gelatin content allowed for a larger
force at break. PCL125g50V showed the highest force at break. The degradation study
was performed to visually examine the fibre structure after 30 days in cell culture
medium at 37 °C. The addition of gelatin stabilised the scaffolds and localised fibre
thinning was only seen in pure PCL samples and no visible fibre degradation points
were seen in the fibres of the blended scaffolds. Gelatin loss in all the samples was also
quantified. It was observed that the gelatin loss in PCL100g75V was maximal in the
first 15 days, while the other ratio PCL125g50V tended to retain its gelatin
concentration without much loss. This is relevant in terms of cell seeding.
The second hypothesis stated that the obtained hybrid morphology (in terms of
structure and composition) enhances cell infiltration, viability and metabolism. To
verify this, the scaffolds were seeded with two different cell types (a robust NIH 3T3
cell line and sensitive primary amnion derived MSCs) to assess if the change in fibre
diameter and pore size distributions allowed for better cell infiltration. Cells showed a
much higher infiltration depth in the hybrid scaffolds (PCL-gelatin blends with fibre
diameter gradients) with a good viability and metabolic activity. PCL175V in
comparison presented hydrophobicity, lower mechanical stability, faster degradation
and lower biocompatibility.
PCL125g50V especially outperformed all the other scaffolds in terms of all the
characterisation tests. It showed the highest infiltration depth in 15 days, had a good
viability percentage and high metabolic activity by day 7 in both cell types. There was
no obstruction of cell penetration because of low pore size or high gelatin concentration.
Therefore, it appears to be the superior scaffold both in terms of structure and
composition for the purpose of cell cultivation.
The aim of the thesis was to fabricate PCL-gelatin blend electrospun scaffolds
with large fibre diameter gradients for the enhanced infiltration and integration of
seeded cells in soft tissue engineering applications. The experimental results show that
86
Chapter 5: Conclusion
the aim has been fulfilled. The complex process of electrospinning is not yet fully
understood, especially the spinning of blended solutions. This research brings us a step
closer to comprehending the various forces involved during the process and the
interaction between them when the setup orientation is changed. Until now, gravity has
been dismissed as a force too weak to have any real influence on the electrospinning
process. However, we see that this is not the case when other parameters such as the
direction of the electric field and conductivity (because of the addition of
polyelectrolyte elements) are altered.
The characterisation data presented in this thesis helped to better understand
cellular response on hybrid PCL-gelatin scaffolds. Much research has gone into
increasing the pore size of scaffolds for better cellular penetration. But are we
compromising on cell function in the process? This is important because in vivo
outcomes are strongly dependent not only on cell infiltration, but also on cell
attachment and overall metabolic health. Since it is extremely difficult to optimise a
scaffold on every aspect, a tradeoff is necessary. The combination of a truly
heterogeneous microstructure and the presence of a bioactive material such as gelatin
can significantly improve how the cells integrate with the scaffold, as can be seen from
the results presented here.
This research also shows that using a robust cell line to validate a scaffold can
be misleading as these cells do not behave like native cells in the body. It is imperative
to gain as much information as possible about sub optimal growth conditions in vitro in
order to avoid failure in vivo. And this can only be achieved by using sensitive cells (for
example, the aMSCs used in this thesis show that pure PCL is not only less favourable
than the gelatin blends, but also actively hampers cell function).
Pore size and fibre diameter have to be optimised for the particular cell type
used in the application. However, this can prove difficult if a co-culture is necessary.
This problem can somewhat be tackled with a heterogeneously structured scaffold like
that presented in this thesis. It is possible to culture multiple cell types on these
electrospun constructs, which can consequently be applicable for a variety of tissues
with mixed cell populations.
87
Chapter 5: Conclusion
5.2 OUTLOOK
Biological systems are so beautifully and intricately designed that a full
understanding may be far in the future. Creating a perfectly biomimetic scaffold of such
a complicated system is even farther off. In the grand scheme of things, it is a much
harder and more complex task to create a versatile scaffold, superior in all aspects for a
particular intended application. Nevertheless, the economic growth within the tissue
engineering sector has been exponential in the recent past, demonstrating tremendous
potential for the future. Engineered tissue has not only been used in regenerative
medicine but also to serve as enabling technologies for other emerging fields such as
bio-robotics, drug delivery and disease modelling.
There are some key limitations to the present study that need to be overcome in
the future. Electrospinning is labour intensive and time consuming and is very hard to
standardise because of the multitude of controlling parameters. In addition, fluorescent
imaging and analysis was made difficult by the autofluorescence of gelatin.
Histological processing was attempted to visualise the distribution of seeded
cells at the end of the infiltration period. Unfortunately, the scaffolds are temperature
sensitive (PCL melting point is 60 °C and gelatin’s is 40 °C) and easily damaged by
alcohol processing, making it hard to get any usable samples for imaging. However,
cryosectioning was attempted on PCL-gelatin tubes (experiments performed in the
Medical University of Graz, Austria as part of an IP@Leibniz funded exchange) and a
preliminary positive result was obtained (Fig 5.1). More experiments in this direction
will be performed in the future.
Further research into the effect of polyelectrolytes on blend electrospinning has
to be performed, especially with regards to changing setup orientation. Is this effect also
observed with other polyelectrolytes? Is it possible to control it by changing the
concentration of the polyelectrolyte? It would be interesting to see if the microstructure
(fibre diameter and pore size distributions) can somehow be guided by the angle of the
electrospinning setup. For instance, the degree of heterogeneity (required for a
particular application) can be increased by making the system orientation more vertical
and decreased by making the system orientation more horizontal.
88
Appendix
References
(1) Vacanti CA. The history of tissue engineering. J. Cell. Mol. Med. 2006; 10(3): 569-576.
(2) Jean J, Garcia-Pérez ME, Pouliot R. Bioengineered skin: the self-assembly approach. J Tissue Sci Eng 2011; S5: 001.
(3) Chan BP, Leong KW. Scaffolding in tissue engineering: general approaches and tissue-specific considerations. Eur Spine J. 2008; 17: 467-479.
(4) Nguyen D, Hägg DA, et al. Cartilage tissue engineering by the 3D bioprinting of iPS cells in a nanocellulose/alginate bioink. Scientific reports 2017; 7: Article 658.
(5) Atala A, Bauer SB, et al. Tissue-engineered autologous bladders for patients needing cystoplasty. Lancet 2006; 367: 1241-1246.
(6) Vaz CM, van Tuijl S, et al. Design of scaffolds for blood vessel tissue engineering using a multi-layering electrospinning technique. Acta Biomaterialia 2005; 1(5): 575-582.
(7) Russias J, Saiz E, et al. Fabrication and in vitro characterization of three-dimensional organic/inorganic scaffolds by robocasting. J Biomed Mater Res A. 2007; 83(2): 434-45.
(8) Alaminos M, Sánchez-Quevedo MDC, et al. Construction of a complete rabbit cornea substitute using a fibrin-agarose scaffold. Invest. Ophthalmol. Vis. Sci. 2006; 47(8): 3311-3317.
(9) Nakase Y, Hagiwara A, et al. Tissue engineering of small intestinal tissue using collagen sponge scaffolds seeded with smooth muscle cells. Tissue Eng. 2006; 12(2): 403–12.
(10) Yoo JI, Ashkar S, Atala A. Creation of functional kidney structures with excretion of urine-like fluid in vivo. British Journal of Urology 1997; 80(SUPPL. 2).
(11) Petersen TH, Calle EA, et al. Tissue-engineered lungs for in vivo implantation. Science 2010; 329(5991): 538-41.
(12) Shimizu T, Yamato M, et al. Cell sheet engineering for myocardial tissue reconstruction. Biomaterials 2003; 24(13): 2309–2316.
(13) Niknamasl A, Ostad SN, et al. A new approach for pancreatic tissue engineering: human endometrial stem cells encapsulated in fibrin gel can differentiate to pancreatic islet beta-cell. Cell Biol Int 2014; 38: 1174-1182.
(14) Li M, Zhang P, Zhang D. PVDF piezoelectric neural conduit incorporated pre-differentiated adipose-derived stem cells may accelerate the repair of peripheral nerve injury. Med Hypotheses 2018; 114: 55-57.
(15) O'Brien FJ. Biomaterials and scaffolds for tissue engineering. Materials today 2011; 14(3): 88-95.
(16) Hasan A, Memic A, et al. Electrospun scaffolds for tissue engineering of vascular grafts. Acta Biomaterialia 2014; 10: 11-25.
(17) Curtis MW, Russell B. Cardiac tissue engineering. The Journal of cardiovascular nursing 2009; 24(2): 87-92.
(18) Truskey GA. Advancing cardiovascular tissue engineering. F1000Research 2016; 5: F1000 Faculty Rev-1045.
(19) Howk D, Chu TM. Design variables for mechanical properties of bone tissue scaffolds. Biomed Sci Instrum. 2006; 42: 278-283.
91
Appendix
(20) Little CJ, Bawolin NK, Chen X. Mechanical properties of natural cartilage and tissue-engineered constructs. Tissue Engineering Part B Reviews 2011; 17(4): 213-227.
(21) Hutmacher DW. Scaffolds in tissue engineering bone and cartilage. The Biomaterials: Silver Jubilee Compendium, The Best Papers Published in BIOMATERIALS 1980–2004. 2000; 175-189.
(22) Loh QL, Choong C. Three-dimensional scaffolds for tissue engineering applications: role of porosity and pore size. Tissue Eng Part B Rev. 2013; 19(6): 485–502.
(23) Subia B, Kundu J, Kundu SC. Biomaterial scaffold fabrication techniques for potential tissue engineering applications. Tissue Engineering, Daniel Eberli. IntechOpen 2010.
(24) soft tissue. In Merriam-Webster.com. Retrieved 16 Oct, 2018. (https://www.merriam-webster.com/medical/soft tissue).
(25) Pei B, Wang W, et al. Fiber-reinforced scaffolds in soft tissue engineering. Regenerative Biomaterials 2017; 4(4): 257–268.
(26) Baker BM, Chen CS. Deconstructing the third dimension – how 3D culture microenvironments alter cellular cues. Journal of Cell Science 2012; 125: 1–10.
(27) Deluzio TGB, Seifu DG, Mequanint K. 3D scaffolds in tissue engineering and regenerative medicine: beyond structural templates? Pharm. Bioprocess. 2013; 1(3): 267–281.
(28) Stern-Straeter J, Riedel F, et al. Advances in skeletal muscle tissue engineering. in vivo 2007; 21: 435-444.
(29) Bhardwaj N, Chouhan D, Mandal BB. 3D functional scaffolds for skin tissue engineering. Functional 3D Tissue Engineering Scaffolds - Materials, Technologies and Applications 2018; 345-365.
(30) Ning L, Sun H, et al. 3D bioprinting of scaffolds with living Schwann cells for potential nerve tissue engineering applications. Biofabrication. 2018; 10(3): 035014.
(31) Vaz CM, Tuij SV, et al. Design of scaffolds for blood vessel tissue engineering using a multi-layering electrospinning technique. Acta Biomaterialia 2005; 1(5): 575-582.
(32) Priyadarsini S, Nicholas SE, Karamichos D. 3D stacked construct: a novel substitute for corneal tissue engineering. Methods Mol Biol. 2018; 1697: 173-180.
(33) Goh JCH, Sahoo S. Scaffolds for tendon and ligament tissue engineering. Regenerative Medicine and Biomaterials for the Repair of Connective Tissues, Woodhead Publishing Series in Biomaterials 2010; 452-468.
(34) Gao M, Zhang H, et al. Tissue-engineered trachea from a 3D-printed scaffold enhances whole-segment tracheal repair. Scientific Reports 2017; 7: 5246.
(35) Yao R, Zhang R, et al. In vitro angiogenesis of 3D tissue engineered adipose tissue. Journal of Bioactive and Compatible Polymers 2009; 24: 5-24.
(36) Wang X, Ali MS, Lacerda CMR. A three-dimensional collagen-elastin scaffold for heart valve tissue engineering. Bioengineering 2018; 5(3): E69.
(37) Ficai D, Albu MG, et al. Advances in the field of soft tissue engineering: From pure regenerative to integrative solutions. Nanobiomaterials in Soft Tissue Engineering, Applications of Nanobiomaterials 2015; 5: 355-386.
(38) Yuksel E, Choo J, et al. Challenges in Soft Tissue Engineering. Seminars in Plastic Surgery 2005; 19(3): 261-270.
(39) Sell SA, Wolfe PS, et al. The use of natural polymers in tissue engineering: a focus on electrospun extracellular matrix analogues. Polymers 2010; 2: 522-553.
(40) Frantz C, Stewart KM, Weaver VM. The extracellular matrix at a glance. Journal of Cell Science 2010; 123(24): 4195-4200.
92
Appendix
(41) Alberts B, Johnson A, et al. Cell-Cell adhesion. Molecular Biology of the Cell. 4th edition. New York: Garland Science 2002. (Retrieved from https://www.ncbi.nlm.nih.gov/books/NBK26937/ on 24th Oct, 2018).
(42) He W, Yong T, et al. Fabrication and endothelialization of collagen-blended biodegradable polymer nanofibers: potential vascular graft for blood vessel tissue engineering. Tissue Eng. 2005; 11: 1574-1588.
(43) Stitzel JD, Pawlowski KJ, et al. Arterial smooth muscle cell proliferation on a novel biomimicking, biodegradable vascular graft scaffold. J. Biomater. Appl. 2001; 16: 22-33.
(44) Heydarkhan-Hagvall S, Schenke-Layland K, et al. Three-dimensional electrospun ECM-based hybrid scaffolds for cardiovascular tissue engineering. Biomaterials 2008; 29: 2907-2914.
(45) Thomas V, Zhang X, Vohra YK. A Biomimetic tubular scaffold with spatially designed nanofibers of Protein/PDS bio-blends. Biotechnol. Bioeng. 2009; 104: 1025-1033.
(46) Choi JS, Lee SJ, et al. The influence of electrospun aligned poly(epsilon-caprolactone)/collagen nanofiber meshes on the formation of self-aligned skeletal muscle myotubes. Biomaterials 2008; 29: 2899-2906.
(47) Dhandayuthapani B, Yoshida Y, et al. Polymeric scaffolds in tissue engineering application: a review. International Journal of Polymer Science 2011; 2011:19 pages.
(48) Szentivanyi A, Chakradeo T, et al. Electrospun cellular microenvironments: Understanding controlled release and scaffold structure. Advanced Drug Delivery Reviews 2011; 63(4-5): 209-220.
(49) Zernetsch H, Repanas A, et al. Electrospinning and mechanical properties of polymeric fibers using a novel gap-spinning collector. Fibers Polym 2016; 17(7): 1025-1032.
(50) Basu P, Repanas A, et al. PEO–CMC blend nanofibers fabrication by electrospinning for soft tissue engineering applications. Materials Letters 2017; 195: 10-13.
(51) Al Halabi, Behrens P, Glasmacher B. Application of electrospun piezoelectric PVDF scaffolds for nerve regeneration. Biomedical Engineering 2014; 59(S1): 193-197.
(52) Mueller M, Bode M, et al. Electrospun vascular grafts with anti-kinking properties. Current Directions in Biomedical Engineering 2015; 1(1): 524–528.
(53) Gilbert, W. De Magnete Magnetcisque Corporibus, et de Magno Magnete Tellure (On the Magnet, Magnetick Bodies also, and on the Great Magnet the Earth; a new Physiology, demonstrated by many arguments & Experiments). Price DJ, editor. London: The Chiswick Press 1600.
(54) Formhals, A. Process and apparatus for preparing artificial threads. US Patent 1975504. 1934.
(55) Formhals, A. Production of artificial fibers. US Patent 2077373. 1937.
(56) Formhals, A. Artificial fiber construction. US Patent 2109333. 1938.
(57) Formhals, A. Method and apparatus for the production of fibers. US Patent 2116942. 1938.
(58) Formhals, A. Method of producing artificial fibers. US Patent 2158415. 1939.
(59) Formhals, A. Method and apparatus for the production of artificial fibers. US Patent 2158416. 1939.
(60) Formhals, A. Artificial thread and method of producing same. US Patent 2187306. 1940.
(61) Formhals, A. Production of artificial fibers from fiber forming liquids. US Patent 2323025. 1943.
(62) Taylor, G. Disintegration of water drops in an electric field. Proceedings of the Royal Society of London Series A Mathematical and Physical Sciences 1964; 280: 383-97.
93
Appendix
(63) Tucker N, Stanger JJ, et al. The History of the science and technology of electrospinning from 1600 to 1995. Journal of Engineered Fibers and Fabrics 2012; Special Issue: 63-73.
(64) Doshi J, Reneker DH. Electrospinning process and applications of electrospun fibers. Journal of Electrostatics 1995; 35(2-3): 151-160.
(65) Ratner BD, Hoffman AS, et al. Electrospinning fundamentals and applications. Biomaterials Science (Third Edition) An Introduction to Materials in Medicine, Academic Press 2012; 332-339.
(66) Khorshidi S, Solouk A, et al. A review of key challenges of electrospun scaffolds for tissue-engineering applications. J Tissue Eng Regen Med 2016; 10: 715–738.
(67) Canbolat MF, Tang C, et al. Mammalian cell viability in electrospun composite nanofiber structures. Macromol Biosci 2011; 11: 1346–1356.
(68) Weldon CB, Tsui JH, et al. Electrospun drug-eluting sutures for local anesthesia. J Control Release 2012; 161: 903–909.
(69) Nam J, Huang Y, et al. Materials selection and residual solvent retention in biodegradable electrospun fibers. J Appl Polym Sci 2008; 107(3): 1547–1554.
(70) Vaquette C, Cooper-White JJ. A simple method for fabricating 3D multilayered composite scaffolds. Acta Biomater 2012; 7: 3277–3284.
(71) Wu J, Hong Y. Enhancing cell infiltration of electrospun fibrous scaffolds in tissue regeneration. Bioactive Materials 1 2016; 1(1): 56-64.
(72) Bhardwaj N, Kundu SC. Electrospinning: A fascinating fiber fabrication technique. Biotechnology Advances 2010; 28(3): 325-347.
(73) Baker BM, Handorf AM, et al. New Directions in Nanofibrous Scaffolds for Soft Tissue Engineering and Regeneration. Expert Rev Med Devices. 2009; 6(5): 515–532.
(74) Tian F, Hosseinkhani H, et al. Quantitative analysis of cell adhesion on aligned micro- and nanofibers. Journal of Biomedical Materials Research A 2008; 84: 291-299.
(75) Teixeira AI, Abrams GA, et al. Epithelial contact guidance on well- defined micro- and nanostructured substrates. J Cell Sci 2003; 116(Pt 10): 1881–1892.
(76) Andersson AS, Backhed F, et al. Nanoscale features influence epithelial cell morphology and cytokine production. Biomaterials 2003; 24(20): 3427–3436.
(77) Dalby MJ, Childs S, et al. Fibroblast reaction to island topography: changes in cytoskeleton and morphology with time. Biomaterials 2003; 24(6): 927–935.
(78) Xie J, Willerth SM, et al. The differentiation of embryonic stem cells seeded on electrospun nanofibers into neural lineages. Biomaterials 2009; 30(3): 354–362.
(79) Li WJ, Mauck RL, et al. Engineering controllable anisotropy in electrospun biodegradable nanofibrous scaffolds for musculoskeletal tissue engineering. J Biomech 2007; 40(8): 1686–1693.
(80) Baker BM, Mauck RL. The effect of nanofiber alignment on the maturation of engineered meniscus constructs. Biomaterials 2007; 28(11): 1967–1977.
(81) Zander NE, Orlicki JA, Rawlett AM. Electrospun polycaprolactone scaffolds with tailored porosity using two approaches for enhanced cellular infiltration. Mater Sci Mater Med 2013; 24: 179–187.
(82) Rnjak-Kovacina J, Wise SG, et al. Tailoring the porosity and pore size of electrospun synthetic human elastin scaffolds for dermal tissue engineering. Biomaterials 2011; 32: 6729–6736.
(83) Balguid A, Mol A, et al. Tailoring fiber diameter in electrospun poly(ε-caprolactone) scaffolds for optimal cellular infiltration in cardiovascular tissue engineering. Tissue Eng 2009; 15: 437–444.
(84) Zhang Y, Ouyang H, et al. Electrospinning of gelatin fibers and gelatin/pcl composite fibrous scaffolds. J Biomed Mater Res B Appl Biomater 2005; 72(1): 156-165.
94
Appendix
(85) Baker BM, Gee AO, et al. The potential to improve cell infiltration in composite fiber-aligned electrospun scaffolds by the selective removal of sacrificial fibers. Biomaterials 2008; 29(15): 2348-2358.
(86) Phipps MC, Clem WC, et al. Increasing the pore sizes of bone-mimetic electrospun scaffolds comprised of polycaprolactone, collagen I and hydroxyapatite to enhance cell infiltration. Biomaterials 2012; 33: 524–534.
(87) Whited BM, Whitney JR, et al. Pre-osteoblast infiltration and differentiation in highly porous apatite-coated PLLA electrospun scaffolds. Biomaterials 2011; 32(9): 2294–2304.
(88) Nam J, Huang Y, et al. Improved cellular infiltration in electrospun fiber via engineered porosity. Tissue Eng 2007; 13(9): 2249-57.
(89) Leong MF, Chan WY, Chian KS. Cryogenic electrospinning: proposed mechanism, process parameters and its use in engineering of bilayered tissue structures. Nanomedicine (Lond) 2013; 8(4): 555-66.
(90) Jiang J, Carlson MA, et al. Expanding two- dimensional electrospun nanofiber membranes in the third dimension by a modified gas-foaming technique. ACS Biomater Sci Eng 2015; 1(10): 991–1001.
(91) Hwang PTJ, Murdock K, et al. Poly(ɛ-caprolactone)/gelatin composite electrospun scaffolds with porous crater-like structures for tissue engineering. J. Biomed Mater Res A 2016; 104(4): 1017–1029.
(92) Muerza-Cascante ML, Haylock D, et al. Melt electrospinning and its technologization in tissue engineering. Tissue Engineering Part B: Reviews 2015; 21(2): 187-202.
(93) Ki CS, Park SY, et al. Development of 3-d nanofibrous fibroin scaffold with high porosity by electrospinning: Implications for bone regeneration. Biotechnol Lett 2008; 30(3): 405-10.
(94) Blakeney BA, Tambralli A, et al. Cell infiltration and growth in a low density, uncompressed three-dimensional electrospun nanofibrous scaffold. Biomaterials 2011; 32(6): 1583-1590.
(95) Zhu X, Cui W, et al. Electrospun fibrous mats with high porosity as potential scaffolds for skin tissue engineering. Biomacromolecules 2008; 9(7): 1795-1801.
(96) Park SH, Kim TG, et al. Development of dual scale scaffolds via direct polymer melt deposition and electrospinning for applications in tissue regeneration. Acta Biomater 2008; 4(5): 1198-1207.
(97) Joshi VS, Lei NY, et al. Macroporosity enhances vascularization of electrospun scaffolds. J Surg Res 2013; 183: 18–26.
(98) Stankus JJ, Guan J, et al. Microintegrating smooth muscle cells into a biodegradable, elastomeric fiber matrix. Biomaterials 2006; 27(5): 735-744.
(99) Li D, Frey MW, Joo YL. Characterization of nanofibrous membranes with capillary flow porometry. J Membr Sci 2006; 286: 104–114.
(100) Eichhorn SJ, Sampson WW. Relationships between specific surface area and pore size in electrospun polymer fibre networks. Journal of the Royal Society Interface 2010; 7(45): 641-649.
(101) Kidoaki S, Kwon IK, Matsuda T. Mesoscopic spatial designs of nano- and microfiber meshes for tissue-engineering matrix and scaffold based on newly devised multilayering and mixing electrospinning techniques. Biomaterials 2005; 26: 37–46.
(102) Ju YM, Choi JS, et al. Bilayered scaffold for engineering cellularized blood vessels. Biomaterials 2010; 31: 4313–4321.
(103) Pham QP, Sharma U, Mikos AG. Electrospun poly(e-caprolactone) microfiber and multilayer nanofiber/microfiber scaffolds: characterization of scaffolds and measurement of cellular infiltration. Biomacromolecules 2006; 7(10): 2796-2805.
95
Appendix
(104) Soliman S, Pagliari S, et al. Multiscale three-dimensional scaffolds for soft tissue engineering via multimodal electrospinning. Acta Biomaterialia 2010; 6: 1227-1237.
(105) Gunatillake PA, Adhikari R. Biodegradable synthetic polymers for tissue engineering. European Cells and Materials 2003; 5: 1-16.
(106) Jagur-Grodzinski J. Polymers for tissue engineering, medical devices, and regenerative medicine. Concise general review of recent studies. Polym. Adv. Technol. 2006; 17: 395–418.
(107) Gigliobianco G, Chong CK, MacNeil S. Simple surface coating of electrospun poly-L-lactic acid scaffolds to induce angiogenesis. Journal of Biomaterials Applications 2015; 30(1): 50–60.
(108) Safaeijavan R, Soleimani M, et al. Biological behavior study of gelatin coated PCL nanofiberous electrospun scaffolds using fibroblasts. Journal of Paramedical Sciences 2014; 5(1): 67-73.
(109) Truong YB, Glattauera V, et al. Collagen-based layer-by-layer coating on electrospun polymer scaffolds. Biomaterials 2012; 33(36): 9198-9204.
(110) Bacakova M, Musilkova J, et al. The potential applications of fibrin-coated electrospun polylactide nanofibers in skin tissue engineering. International Journal of Nanomedicine 2016; 11: 771-789.
(111) Hinderer S, Sudrow K, et al. Surface functionalization of electrospun scaffolds using recombinant human decorin attracts circulating endothelial progenitor cells. Scientific Reports 2018; 8: article number 110.
(112) Liu B, Xu F, et al. Electrospun PLLA fibers coated with chitosan/heparin for scaffold of vascular tissue engineering. Surface & Coatings Technology 2013; 228: S568–S573.
(113) Hackett JM, TT Dang, et al. Electrospun Biocomposite polycaprolactone/collagen tubes as scaffolds for neural stem cell differentiation. Materials 2010; 3: 3714-3728.
(114) Daelemans L, Steyaert I, et al. Nanostructured hydrogels by blend electrospinning of polycaprolactone/gelatin nanofibers. Nanomaterials 2018; 8(7): 551.
(115) Suresh S, Gryshkov O, Glasmacher B. Impact of setup orientation on blend electrospinning of poly-ε-caprolactone-gelatin scaffolds for vascular tissue engineering. The International Journal of Artificial Organs 2018; 41(11): 801–810. (Images and tables reused (with and without modification) with permission form SAGE publishing)
(116) Li M, Mondrinos MJ, et al. Electrospun blends of natural and synthetic polymers as scaffolds for tissue engineering. Conf Proc IEEE Eng Med Biol Soc. 2005; 6: 5858-61.
(117) Sell SA, Wolfe PS, et al. Incorporating platelet-rich plasma into electrospun scaffolds for tissue engineering applications. Tissue Eng 2011; 17: 2723–2737.
(118) Lee J, Yoo JJ, et al. The effect of controlled release of PDGF-BB from heparin-conjugated electrospun PCL/gelatin scaffolds on cellular bioactivity and infiltration. Biomaterials 2012; 33: 6709–6720.
(119) Chakraborty S, Liao I, et al. Electrohydrodynamics: A facile technique to fabricate drug delivery systems. Adv Drug Deliv Rev. 2009; 61(12): 1043–1054.
(120) Yang C, Jia Z, et al. Comparisons of fibers properties between vertical and horizontal type electrospinning systems. Annual Report - Conference on Electrical Insulation and Dielectric Phenomena, CEIDP 2009: 204-207.
(121) Rodoplu D, Mutlu M. Effects of electrospinning setup and process parameters on nanofiber morphology intended for the modification of quartz crystal microbalance surfaces. Journal of Engineered Fibers and Fabrics 2012; 7(2): 118-123.
(122) Li Y, Liu J, et al. Optimization of the electrospinning process for core–shell fiber preparation. Journal of Biomaterials and Tissue Engineering 2014; 4(11): 973–980.
96
Appendix
(123) Azimi B, Nourpanah P, et al. Poly (ε-caprolactone) fiber: an overview. Journal of Engineered Fibers and Fabrics 2014; 9(3): 74-90.
(124) Ulery BD, Nair LS, Laurencin CT. Biomedical applications of biodegradable polymers. J Polym Sci B Polym Phys. 2011; 49(12): 832–864.
(125) Woodruff MA, Hutmacher DW. The return of a forgotten polymer—Polycaprolactone in the 21st century. Progress in Polymer Science 2010; 35(10): 1217-1256.
(126) Hutmacher DW. Scaffold design and fabrication technologies for engineering tissues – state of the art and future perspectives. J Biomater Sci Polym Ed 2001; 12: 107–24.
(127) Lam CXF, Savalani MM, et al. Dynamics of in vitro polymer degradation of polycaprolactone-based scaffolds: accelerated versus simulated physiological conditions. Biomed Mater 2008; 3: 1-15.
(128) Sbyrnes321. Own work 2012. (Retrieved from https://commons.wikimedia.org/w/index.php?curid=18040958 on 25th Oct, 2018). License: CC0 1.0 Universal (https://creativecommons.org/publicdomain/zero/1.0/legalcode)
(129) Chaplin M. Gelatin. Water structure and Science 2012. (Retrieved from http://www1.lsbu.ac.uk/water/gelatin.html on 25th Oct, 2018). License: CC BY-NC-ND 2.0 UK Attribution - Non-Commercial - NoDerivs 2.0 England and Wales (https://creativecommons.org/licenses/by-nc-nd/2.0/uk/legalcode)
(130) Sivabalan A, Harihara subramani R, et al. Synthesis and characterization of poly (ε-caprolactone): a comparative study. International Journal of Scientific Research Engineering & Technology (IJSRET) , ICRTIET Conference Proceeding 2014: 9-14.
(131) Elzoghby AO. Gelatin-based nanoparticles as drug and gene delivery systems: Reviewing three decades of research. Journal of Controlled Release 2013; 172: 1075–1091.
(132) Hoque1 ME, Nuge T, et al. Gelatin based scaffolds for tissue engineering - a review. Polymers Research Journal 2015; 9(1): 15-32.
(133) Wang H, Boerman OC, et al. Comparison of micro- vs. nanostructured colloidal gelatin gels for sustained delivery of osteogenic proteins: Bone morphogenetic protein-2 and alkaline phosphatase. Biomaterials 2012; 33: 8695-8703.
(134) Gorgieva S, Kokol V. Collagen- vs. Gelatine-based biomaterials and their biocompatibility: review and perspectives. In: Biomaterials Applications for Nanomedicine. Prof. Rosario Pignatello (Ed.), InTech 2011; 17-52.
(135) Li W-J, Shanti RM, Tuan RS. Electrospinning technology for nanofibrous scaffolds in tissue engineering. In: Nanotechnologies for the Life Sciences Vol. 9: Tissue, Cell and Organ Engineering' (ed.: Kumar C.) Wiley-VCH Verlag GmbH & Co. KGaA, Germany, 2006; 135-187.
(136) Fodor WL. Tissue engineering and cell based therapies, from the bench to the clinic: the potential to replace, repair and regenerate. Reprod Biol Endocrinol. 2003; 1: 102.
(137) Kaur G, Dufour JM. Cell lines: Valuable tools or useless artifacts. Spermatogenesis 2012; 2(1): 1-5.
(138) Pogozhykh O, Pogozhykh D, et al. Molecular and cellular characteristics of human and non-human primate multipotent stromal cells from the amnion and bone marrow during long term culture. Stem Cell Research & Therapy 2015; 6(150): 15 pages.
(139) Mukhopadhyay R. Generating the NIH 3T3 cell line, the oncogene hypothesis and horses. ASBMB Today, Generations 2015. (Retrieved from http://www.asbmb.org/asbmbtoday/201501/Generations/3T3/ on 22nd Oct, 2018).
(140) NIH 3T3 Cell Line (mouse embryonic fibroblasts) NIH3T3 General Information. Retrieved from http://www.nih3t3.com/ on 22nd Oct, 2018.
97
Appendix
(141) Greig RG, Koestler TP, et al. Tumorigenic and metastatic properties of “normal” and ras-transfected NIH/3T3 cells. Proceedings of the National Academy of Sciences of the United States of America. 1985; 82(11): 3698-3701.
(142) Gryshkov O, Pogozhykh D, et al. Encapsulating non-human primate multipotent stromal cells in alginate via high voltage for cell-based therapies and cryopreservation. PLoS One 2014; 9(9): e107911.
(143) Chong EJ, Phan TT, et al. Evaluation of electrospun PCL/gelatin nanofibrous scaffold for wound healing and layered dermal reconstitution. Acta Biomaterialia 2007; 3(3): 321-330.
(144) Ghasemi-Mobarakeh L, Prabhakaran MP, et al. Electrospun poly(ɛ-caprolactone)/gelatin nanofibrous scaffolds for nerve tissue engineering. Biomaterials 2008; 29(34): 4532-4539.
(145) Zheng R, Duan H, et al. The influence of Gelatin/PCL ratio and 3-D construct shape of electrospun membranes on cartilage regeneration. Biomaterials 2014; 35(1): 152-164.
(146) Fu W, Liu Z, et al. Electrospun gelatin/PCL and collagen/PLCL scaffolds for vascular tissue engineering. International Journal of Nanomedicine 2014; 9(1): 2335-2344.
(147) Kai D, Prabhakaran MP, et al. Guided orientation of cardiomyocytes on electrospun aligned nanofibers for cardiac tissue engineering. J. Biomed. Mater. Res., 2011; 98B: 379-386.
(148) Feng B, Tu H, et al. Acetic-acid-mediated miscibility toward electrospinning homogeneous composite nanofibers of GT/PCL. Biomacromolecules 2012; 13(12): 3917–3925.
(149) McCarthy J, Bhattacharya S, et al. Composition and stress analysis in Si structures using micro-Raman spectroscopy. Scanning 2004; 26: 235–239.
(150) Class II Biological Safety Cabinets and Laminar Airflow benches. Thermo Scientific. (Retrieved from https://www.thermofisher.com/de/de/home/life-science/lab-equipment/biological-safety-cabinets-clean-benches.html on 14th Nov, 2018.)
(151) Dong Y, Yong T, et al. Degradation of electrospun nanofiber scaffold by short wave length ultraviolet radiation treatment and its potential applications in tissue engineering. Tissue Eng. Part A 2008; 14: 1321.
(152) Phalloidin, Fluorescein Isothiocyanate Labeled. Sigma Aldrich (Retrieved from https://www.sigmaaldrich.com/catalog/product/sigma/p5282?lang=de®ion=DE on 12th Nov, 2018.)
(153) bisBenzimide H 33342 trihydrochloride. Sigma Aldrich (Retrieved from https://www.sigmaaldrich.com/catalog/product/sial/b2261?lang=de®ion=DE on 12th Nov, 2018.)
(154) ThermoFisher Scientific., Live/Dead Viability/Cytotoxicity Kit for mammalian cells, Germany, (Retrieved from: http://www.thermofisher.com/order/catalog/product/L3224.)
(155) Riss TL, Moravec RA, et al. Cell Viability Assays. In: Sittampalam GS, Coussens NP, Brimacombe K, et al., editors. Assay Guidance Manual [Internet]. Bethesda (MD): Eli Lilly & Company and the National Center for Advancing Translational Sciences; 2004-. 2013 [Updated 2016 Jul 1] (Available from: https://www.ncbi.nlm.nih.gov/books/NBK144065/)
(156) Müller M. Faserbasierte abbaubare kardiovaskuläre Gefaßprothesen: Entwicklung, Herstellung und Prüfung (PhD thesis). Institut für Mehrphasenprozesse 2018; Band 1.
(157) Frushour BG, Koenig JL. Raman scattering of Collagen, Gelatin, and Elastin. Biopolymers 1975; 14: 379-391.
(158) Wesełucha-Birczyńska A, Świętek M, et al. Raman spectroscopy and the material study of nanocomposite membranes from poly(ε-caprolactone) with biocompatibility testing in osteoblast-like cells. Analyst 2015; 140: 2311-2320.
(159) Agrahari V, Agrahari V, et al. Electrospun nanofibers in drug delivery: fabrication, advances, and biomedical applications (Chapter 9). Emerging Nanotechnologies for Diagnostics, Drug Delivery and Medical Devices, Micro and Nano Technologies 2017; 189-215.
98
Appendix
(160) Son WK, Youk JH, et al. The effects of solution properties and polyelectrolyte on electrospinning of ultrafine poly(ethylene oxide) fibers. Polymer 2004; 45(9): 2959-2966.
(161) Taylor SG. Electrically driven jets. Proc. R. Soc. Lond. A 1969; 313: 453-475. License: The Royal Society Academic Institution Single Site License (“the Site Licence”) based on the Model Licence Version 4.0, Oct 6, 2009 for subscription articles (https://royalsociety.org/about-us/terms-conditions-policies/publishing-terms-conditions/academic-single-site/). Access granted through TIB subscription.
(162) Reneker DH, Yarin AL. Electrospinning jets and polymer nanofibers. Polymer 2008; 49(10): 2387-2425.
(163) Pfeifer RJ, Hendricks Jr CD. Parametric studies of electrohydrodynamic spraying. AIAA Journal 1968; 6(3): 496 - 502.
(164) Horning DW, Hendricks CD. Study of an electrically driven jet. J. Appl. Phys. 1979; 50(4): 2614-2617.
(165) Stanger J, Staiger MP, et al. Effect of charge density on the Taylor cone in electrospinning. Solid State Phenomena 2009; 151: 54-59.
(166) Nelson MT, Johnson J, Lannutti JJ. Media-based effects on the hydrolytic degradation and crystallization of electrospun synthetic-biologic blends. Mater Sci: Mater Med 2014; 25(2): 297-309.
(167) Zhou B, Xu F, et al. Strain rate sensitivity of skin tissue under thermomechanical loading. Journal of the Royal Society 2010; 368: 679-690.
(168) Ankerson J, Birkbeck AE, et al. Puncture resistance and tensile strength of skin simulants. Journal of Engineering in Medicine 1999; 213: 493-501.
(169) Innocenti B, Larrieu JC, et al. Automatic characterization of soft tissues material properties during mechanical tests. Muscles Ligaments Tendons J. 2018; 7(4): 529-537.
(170) Raghavan BK, Coffin DW. Control of inter-fiber fusing for nanofiber webs via electrospinning. J Eng Fiber Fabr 2011; 6(4): 1–5.
(171) Argento G. Biomechanical design of an electrospun scaffold for in situ cardiovascular tissue engineering (PhD Thesis). Technische Universiteit Eindhoven, Eindhoven, 2014.
(172) Kancheva M, Toncheva A, et al. Enhancing the mechanical properties of electrospun polyester mats by heat treatment. Express Polym Lett 2015; 9(1): 49–65.
(173) Kotula AP, Snyder CR, Migler KB. Determining conformational order and crystallinity in polycaprolactone via Raman spectroscopy. Polymer (Guildf) 2017; 117: 1-10.
(174) Heimowska A, Morawska M, Bocho-Janiszewska A. Biodegradation of poly(ε-caprolactone) in natural water environments. Polish Journal of Chemical Technology 2017; 19(1): 120-126.
(175) Chen M, Patra PK, et al. Role of electrospun fibre diameter and corresponding specific surface area (SSA) on cell attachment. J Tissue Eng Regen Med 2009; 3: 269-279.
(176) Kim J, Lee Y, et al. Human amniotic fluid-derived stem cells have characteristics of multipotent stem cells. Cell Proliferation 2007; 40: 75-90.
(177) Jin HJ, Bae YK, et al. Comparative analysis of human mesenchymal stem cells from bone marrow, adipose tissue, and umbilical cord blood as sources of cell therapy. Int. J. Mol. Sci. 2013; 14: 17986-18001.
(178) Heo JS, Choi Y, et al. Comparison of molecular profiles of human mesenchymal stem cells derived from bone marrow, umbilical cord blood, placenta and adipose tissue. Int J Mol Med. 2015; 37(1):115-25.
99
Appendix
(179) Åsa Ekblad. Comparison of Mesenchymal Stem Cells derived from amniotic fluid and umbilical cord (Master thesis). Department of Clinical Sciences, Intervention and Technology, Karolinska Institutet, Sweden. 2011.
(180) Yao R, He J, Meng G, et ak. Electrospun PCL/Gelatin composite fibrous scaffolds: mechanical properties and cellular responses. Journal of Biomaterials Science, Polymer Edition 2016; 27(9): 824-838.
(181) Sun M, Chi G, et al. Effects of matrix stiffness on the morphology, adhesion, proliferation and osteogenic differentiation of mesenchymal stem cells. Int J Med Sci. 2018; 15(3): 257-268.
(182) Leong, MF, Rasheed MZ, et al. In vitro cell infiltration and in vivo cell infiltration and vascularization in a fibrous, highly porous poly(D,L-lactide) scaffold fabricated by cryogenic electrospinning technique. J. Biomed. Mater. Res. 2009; 91A: 231-240.
(183) Wu J, Liu S, et al. Electrospun nanoyarn scaffold and its application in tissue engineering. Mater. Lett. 2012; 89: 146-149.
(184) Lee JB, Jeong SI, et al. Highly porous electrospun nanofibers enhanced by ultrasonication for improved cellular infiltration. Tissue Eng Part A. 2011; 17(21-22): 2695-702.
100
Appendix
Supplementary data
S.1 Thickness optimisation of electrospun scaffolds
The spinning time for the 150 µm thick scaffolds was extrapolated from the
thickness optimisation graph. Since this would only be an estimate, samples were spun
for the suggested time period and then measured again for thickness. Any variation in
empirical thickness to the estimated thickness was reduced by simply adjusting the
spinning time. These samples were measured again to ensure that the spinning time
indeed produced the required scaffold thickness.
Name of the solution Optimised duration of electrospinning (min)
50 µm thickness 150 µm thickness
PCL175V 25 135
PCL125g50H 110 332
PCL125g50V 120 232
PCL100g75V 85 172
101
Appendix
S.2 Mechanical testing Force-Elongation graphs (all sets)
102
Appendix
S.3 Raman spectra - PCL pellets, gelatin powder
103
Appendix
S.4 Degradation study - Raman spectra (fingerprint region)
104
Appendix
105
Appendix
Publication list
1) Suresh S, Gryshkov O, Glasmacher B. Impact of setup orientation on blend
electrospinning of poly-ε-caprolactone-gelatin scaffolds for vascular tissue
engineering. The International Journal of Artificial Organs 2018; 41(11): 801–810.
2) Suresh S, Balamuruganandam MS, Kidwai SM, Gryshkov O, Glasmacher B.
Improving cell infiltration in electrospun poly(ε-caprlactone)-gelatin scaffolds
through fibre diameter manipulation: a comparative analysis with amnion derived
multopotent stromal cells and 3T3 cell line. Materials Science and Tissue
Engineering C [under review as of 21st February, 2019].
108
Appendix
Statement of Contribution
I, Sinduja Suresh, hereby confirm that I have made the following contributions to this
thesis:
1) Supervised two relevant master thesis projects to completion.
• Investigation of effect of system orientation on properties of electrospun PCL
gelatin scaffolds intended for cardiovascular application.
(Mythili Shree Balamuruganandam, Vellore Institute of Technology, India,
July 2017)
Relevant topics: Electrospinning, mechanical study
• A comparative cell response study on structurally enhanced Poly-ε-
caprolactone (PCL)-Gelatin electrospun scaffolds for cardiovascular
applications.
(Sarah Manaal Kidwai, University of Göttingen, Germany, October 2018)
Relevant topics: aMSC infiltration, viability and metabolic activity
2) Supervised one relevant lab rotation
• Assessment of metabolic activity of NIH 3T3 fibroblasts seeded on PCL-
gelatin electrospun scaffolds intended for cardiovascular applications.
(Sarah Manaal Kidwai, University of Göttingen, Germany, April 2018)
Relevant topics: 3T3 metabolic activity
3) Supervised one relevant exchange project funded by IP@Leibniz.
• Investigation of infiltration of 3T3 fibroblasts into electrospun PCL-gelatin
scaffolds intended for cardiovascular applications.
(Mythili Shree Balamuruganandam, November 2017)
Relevant topics: 3T3 infiltration, viability
111
Appendix
(Note for points 1 to 3: I conceptualised and supervised the optimisation of protocols.
All experiments were performed by me and the student together in order to accomplish
more in a short amount of time. Optimised protocols and results were verified and
approved by Dr Oleksandr Gryshkov.)
4) Supervised one relevant student project to completion.
Charakterisierung von tubular PCL-Gelatine-Elektrospinn Gerüststrukturen
(Mugtaba Azim, Leibniz University Hannover, Germany, July 2018)
(This project was done to make sure that the electrospinning machine did not
influence the PCL-gelatin microstructure.)
5) Supervised four internships to completion (data not included in this thesis)
• Construction of a coaxial needle and electrohydrodynamic unit
(Nicholas Feringa, Michigan State University, August 2016)
• Parametric optimisation of electrosprayed alginate beads
(Michaela Keck, University of Nebraska-Lincoln, August 2016)
• Fabrication and evaluation of alginate Volvox spheres
(Sarah Cushman, Northeastern University, December 2016)
• Optimisation of cell-seeding on electrospun scaffolds using micro extrusion
printing
(Cing Hawm Lian, University of Nebraska-Lincoln, August 2017)
6) Prepared and submitted the abstracts and associated poster/talk for all the
conference contributions (refer CV).
7) Measurements of solution density and viscosity. Conductivity measurements were
done by Benedikt Voß (intern).
8) Analysis of Raman data (actual spectra acquired by Dr Oleksandr Gryshkov
because only authorised personnel are allowed to use the machine).
112
Appendix
9) FTIR spectroscopy (with instruction from Prof. Dr. Ir. Willem F. Wolkers).
10) Designed the scaffold cutters on SolidWorks (cutters manufactured in the IMP
workshop).
11) Put forward the hypothesis explaining the phenomenon of changing
microstructure in PCL-gelatin blend electrospun scaffolds in the vertical
orientation above a critical concentration of gelatin.
12) Did an exchange funded by IP@Leibniz in the Medical University of Graz (refer
CV) under the supervision of Dr Dagmar Pfeiffer. Tubular PCL-gelatin grafts
were seeded with human endothelial and smooth muscle cells in bioreactors. A
protocol for histological processing (either paraffin or cryo) of such temperature-
sensitive scaffolds was developed. The tubular scaffolds used in this project were
electrospun by me at the IMP using the optimised protocol developed during the
student project of Mugtaba Azim.
13) Performed histological staining on flat mats in the Niedersächsisches Zentrum für
Biomedizintechnik, Implantatforschung und Entwicklung (NIFE) with Ms
Annemarie Beck to identify appropriate stains for PCL-gelatin scaffolds (data not
included in this thesis).
14) Attended conferences: ESAO & yESAO 2016 (Warsaw, Poland), yESAO 2017
(Vienna, Austria) and ESAO & yESAO 2018 (Madrid, Spain).
15) Published a first author paper. The second first author paper has been submitted.
(refer publication list)
113