UU student #: 5582733
QUT student #: n9827102
Iris OTTO
Dr.Phong TRAN
Dr.Ir.Jos MALDA
D/Prof.Dietmar W. HUTMACHER
UMC – Utrecht University, Netherlands
IHBI - Queensland University of Technology, Australia
01/09/2015 – 31/08/2017
SC80 Master of Applied Science
MSc Biofabrication
Final Thesis
“Biofabrication: tools for new therapeutics in
regenerative medicine and drug delivery.” Submitted in fulfilment of the requirement for the degree of SC80 Master of Applied Science
Science and Engineering Faculty
Quentin Clément PEIFFER
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STATEMENT OF ORIGINAL AUTHORSHIP
The work contained in this thesis undertaken between QUT and Utrecht University has not
been previously submitted to meet requirements for an award at these or any other higher
education institution. To the best of my knowledge and belief, the thesis contains no material
previously published or written by another person except where due reference is made.
QUT Verified Signature
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FOREWORD
This Master thesis is the result of the collaboration between the Utrecht University (UU) and
the Queensland University of Technology (QUT). As such, this document is divided in two
independent section: the first part “Biofabrication of an auricular cartilage” is presenting the
work carried out at UU while the second part “ Microporous polycaprolactone scaffolds for drug
delivery” is presenting the work carried out at QUT. As such, each part features its own ab-
stract, keywords, abbreviation, acknowledgement, table of content and bibliography.
Biofabrication of an auricular cartilage………………………………p3 - 43
Microporous polycaprolactone scaffolds for drug delivery…………p44-78
P a g e 3 | 78
Biofabrication of an auricular
cartilage implant
Quentin Clément PEIFFER
UU student #: 5582733
MSc Biofabrication
Minor Research Project
RMC Utrecht
09/11/2015 – 10/06/2016
Daily supervisor: Dr.Phong Tran
Examiner: Dr.Ir.J.Malda
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LAYMAN’S SUMMARY
Facial malformations like ear loss due to cancer, burns, trauma or even birth defects can heav-
ily affect the relationship between an individual and their relatives or society. Current treat-
ments present severe drawbacks with highly variable aesthetic results. With the capacity to
combine different materials in a precise manner, 3D printers appeared in biomedical sciences
during the past few years as novel tools able to bring new solutions, such as scaffolds, that
can overcome all previous treatments. A scaffold in tissue engineering is a construct made of
a material compatible with the human body to repair damaged tissues. Yet, to provide these
new clinical solutions, 3D printers require considerable work for research and optimization. In
this work, the deposition of two materials is studied to combine them in an ear shape cartilage.
The first material is a thermoplastic that provides mechanical strength and consequently, the
printed scaffolds is not destroyed after grafting. The second material is a hydrogel; a hydrogel
is a gel able to absorb a high quantity of water and therefore, provides conditions for cells to
proliferate. To combine these two materials, it is first necessary to optimize the deposition of
each of them individually. The first part of this work is to study the potential of a new tool for
the deposition of the thermoplastic material. Since the printing process can kill the cells, the
second step of this work is to review the literature to predict how to deposit the hydrogel while
preserving the highest cell survival. The thermoplastic and hydrogel laden with cells were com-
bined in square constructs and analyzed in the third step of this work. The final phase of this
work focuses on the computer work related to the control of the printer, to assess which soft-
ware would be the most useful for carrying out the printing of two materials in an ear shape.
ABSTRACT
Microtia or ear loss are facial malformations for which no current treatments are perfectly
adapted. Additive manufacturing is a growing field and is expected to provide medical applica-
tions in the near future, especially by the creation of intricate scaffolds. This study explores the
co-manufacturing of hybrid PCL/gelMA scaffolds, specifically for ear cartilage engineering.
This research with a step-by-step approach aims to present the different challenges that
represent co-manufacturing and how they could be overcome. This includes a description of
Fused-deposition modeling (FDM) and Pressure-Assisted Bioprinting (PAB) with attention
given to the preservation of cell-viability. If the combination of FDM and PAB is not a technical
challenge, this work illustrates the importance of characterizing materials rheological proper-
ties to have control over the fabrication process. Therefore, after experimenting and backed
by literature, it appears that spraying cell with low inlet-pressure is the approach that preserves
the highest cell viability when dispensing a cell-laden hydrogel. At last, this work points out the
importance of considering the computer science behind additive manufacturing, and which
otherwise can rapidly become a limitation for tools capacities
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ABBREVIATIONS 3D: Three Dimensional
CAD: Computer-Aided Design
DMEM: Dulbecco's Modified Eagle Medium
ECM: Extracellular Matrix
EXC: Experimental Cartridge
FDA: Food and Drug Administration
GelMA: Gelatin methacryloyl
HMI: Human Machine Interface
PBS: Phosphate-buffered saline
PCL: Polycaprolactone
RGD: Arginylglycylaspartic acid
STL: Standard Tessellation Language
ACKNOWLDEGMENT Jos Malda
Pedro da Costa
Iris Otto
Kim van Dorenmalen
Riccardo Levato
Maarten Blokzjil
Sarah-Sophia Carter
Madeline Hintz
Noël Dautzenberg
KEYWORDS
PCL; Auricular cartilage, GelMA, Dual printing ; 3D printing ; Biofabrication;
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TABLE OF CONTENTS
ABSTRACT ............................................................................................................................................ 4
ABBREVIATIONS ................................................................................................................................. 5
ACKNOWLDEGMENT ......................................................................................................................... 5
KEYWORDS .......................................................................................................................................... 5
1. INTRODUCTION .............................................................................................................................. 7
1.1 Materials .............................................................................................................. 7
1.2 The 3D printing approach ..................................................................................... 8
2. CHAPTER 1: PCL PRINTING ...................................................................................................... 11
2.1 Introduction .........................................................................................................11
2.2 Material and methods..........................................................................................11
2.3 Results ................................................................................................................13
2.4 Discussion ..........................................................................................................17
2.5 Conclusion ..........................................................................................................18
3. Chapter 2: Spraying vs Deposition .............................................................................................. 18
3.1 Introduction .........................................................................................................18
3.2 Conclusion ..........................................................................................................21
4. Chapter 3: Cell viability and printing ............................................................................................ 21
4.1 Introduction .........................................................................................................21
4.2 Material and method ...........................................................................................22
4.4 Troubleshooting ..................................................................................................29
4.5 Discussion ..........................................................................................................30
4.6 Conclusion and Further Experiments ..................................................................33
5. Chapter 4: Auricular shape and biofabrication .......................................................................... 33
5.1 Introduction .........................................................................................................33
5.2 Material and method ...........................................................................................34
5.3 Results and troubleshooting ................................................................................35
5.4 Discussion and Conclusion .................................................................................36
6. Chapter 5: General conclusion and prospective work .............................................................. 36
REFERENCES .................................................................................................................................... 40
ANNEX 1 .............................................................................................................................................. 41
ANNEX 2 .............................................................................................................................................. 42
ANNEX 3 .............................................................................................................................................. 43
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1. INTRODUCTION
Facial malformations hinder the relationship between a patient and society, resulting in a social
and psychological burden, that when surgically treated can highly improve psychosocial as-
pects and consequently the quality of life of the patient. Auricular malformations as a result of
congenital anomalies (microtia), cancer, burns or even trauma, are part of these facial malfor-
mations. Current treatments options for auricular malformation include ear prostheses, syn-
thetic implants and auricular reconstruction using skin flaps or autologous rib cartilage. Be-
cause of the complex three-dimensional (3D) shape of the auricle, auricular reconstruction with
autologous costal cartilage is a challenging procedure with a highly variable aesthetic outcome;
not to mention the significant operating time and donor site morbidity. In response, efforts have
been made towards creating pre-fabricated synthetic auricular implants (Medpor®). Medpor®
appears to be a good solution, even though as a foreign body it can potentially lead to implant
exposure or infection risks. However, the great majority of plastic surgeons prefers the use of
autologous cartilage frameworks, that is the current gold-standard over synthetic implants.
The convergence of regenerative medicine and biofabrication brings new alternatives that
would overcome limitations associated with current treatments such as donor site morbidity
while improving aesthetic and functional outcomes. It allows the possibility to engineer func-
tional cartilage using patient-derived or donor cells, and to create custom-designed cell-laden
implants with intricate architectures and complex shapes.
Biofabrication is particularly interesting since it offers the opportunity to combine different cell
types and materials to produce the ideal auricular scaffold. In addition, it may someday reach
a higher level of complexity, by incorporating fatty tissue or perichondrium. To be successful
the ideal auricular engineered scaffold should:
- Be strong enough to withstand the contractive forces of the skin and durably maintain
the same shape than the contralateral auricle.
- Incorporate autologous chondrocytes or stem cells that will recreate a matrix with the
natural elastic bending properties of the auricle.
- Be slowly degradable while new cartilaginous matrix replaces it, maintaining its original
shape.
Therefore, the project was based on I. Otto’s work [1] and the knowledge from literature.
1.1 Materials
1.1.1 Soft hydrogel: GelMA
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Chondrocytes typically thrive in a soft hydrogel that allows unimpeded nutrient diffusion and
provides a homogenous microenvironment promoting cellular migration, proliferation, and dif-
ferentiation, and to this purpose, GelMA was used, a water-soluble protein that mimics the
Extracellular Matrix (ECM). Biodegradable, biocompatible and able to form hydrogels, many
features make GelMA a good candidate in biomedical science and these are documented in
the literature. Its functionalization with unsaturated methacryloyl, combined with a photoinitiator
and under exposition of UV-light enable the GelMA to form covalently cross-linked hydrogels,
that can be cultured with encapsulated cells.
1.1.2 Stiff polymer: PCL
To withstand the contractive forces of the skin and durably maintain shape, the scaffold will
need to have a high degree of stiffness. A soft hydrogel such as GelMA, even if cross-linked,
cannot by itself reach the degree of stiffness required. Therefore, a stiff polymer, in our case
polycaprolactone (PCL), will be deposited with GelMA to enhance the mechanical properties
of thereof. PCL is a biodegradable polymer, non-toxic, with a broad miscibility, and a mechan-
ical compatibility with many polymers. Furthermore, it provides adhesion to a broad spectrum
of substrates, can be modified to create microporous fibers, or be grafted with the cell adhesion
site (such as RGD). PCL is approved by the FDA for specific biomedical applications and
widely used in research, especially in 3D printing where its mechanical properties and low-
melting-point make it an ideal printable polymer.
These materials are very common in the regenerative medicine landscape, especially in
biofabrication where their features are highly appreciated. The next section below provides a
more technical description of the project.
1.2 The 3D printing approach
1.2.1 Biofabrication approach
A cell-laden GelMA hydrogel solution is deposited between strips of PCL (Fig. 1) with robotic
dispensing technology using layer-by-layer deposition according to a computer-aided design.
This leads to the creation of custom hybrid constructs that combines the cell compatibility of
GelMA and the stiffness of PCL. This approach is already described in the literature and has
been proved viable [2], the thermal requirement for printing PCL was shown to be compatible
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with cell viability in the hydrogel [3]. Furthermore, the real potential of this approach lies in its
versatility, by combining different cells, materials or architecture all kind of applications are
possible [4].
1.2.2 Printers
Several tools are available in the laboratory, depending on the system the biofabrication pro-
cedure will be different slightly different, more details can be found in Chapter 4. In the labor-
atory, two robotic dispensing system is available, the Bioscaffolder and the 3D Discovery.
First the Bioscaffolder (SYS+ENG) (Fig 2 A): The Bioscaffolder is a 3-axis dispensing system
with an automatic tool change function, controlled by a Human Machine Interface (HMI) called
PrimCAM. PrimCAM allows the user to import Standard Tessellation Language (STL) files but
provides poor drawing tools. The Bioscaffolder can possess up to 5 dispense heads allowing
Fig.1. Schematic overview of the hybrid bioprinting process. A three-dimensional design is
translated to a deposition protocol which alternates steps of printing polymer and cell-laden
hydrogels to yield hybrid constructs. Source : [2]
Fig.2 (a) Image of the Bioscaffolder printer and its two printheads (b), PCL auger screw
pump on the left and syringe piston on the right. (c) Visual representation of the missing
heating piece.
(a) (b)
à
(c)
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the printing of a wide range of material from medium to high viscosity biomaterials (biopoly-
mers) to low and medium viscosity biomaterials such as hydrogels and silicones. The Bioscaf-
folder has been the first choice for the realization of the project since it was the only system in
the lab, at that time, able to dispense PCL and a GelMA based hydrogel ink in the same struc-
tures. Usually, the GelMA is heated up to 37°C and dispensed by a mechanical piston which
applies pressure over time to a syringe plunger (Fig. 2 b). Unfortunately, the piece ensuring
the heating of the syringe’s nozzle is not available in the lab (Fig. 2 c), consequently, GelMA
is cooling down in the nozzle impeding a proper deposition between the PCL strands. Finally,
after several attempts to overcome this issue, it was decided to change the tool used.
The 3DDiscovery (regenHU) (Fig 3 a & b): Compared to the Bioscaffolder, the 3D discovery
is a pneumatic based system that can host 4 different printheads in addition to a UV-light tool.
All of them work in a coordinated motion since they are connected to the same robotic arm;
thanks to this feature the 3DDiscovery has a work speed significantly higher than the Bioscaf-
folder. The HMI of the 3D Discovery use its own G-code, therefore it can only import .iso files
created in software provided by regenHU, BioCAD and MMconverter (see chapter 4 for more
details). The 3D Discovery can print a wide range of materials due to its 4 types of the
printhead, nevertheless, the printer can only print thermopolymers and highly viscous media
with an HM-300H printhead (Fig. 3 c). However, our lab made the choice to install two DD-
135N and two CF-300H (Fig. 3). In absence of the HM-300H, it is technically impossible to
print highly viscous material, such as PCL, with the 3Ddiscovery, and it was the main reason
Fig. 3. Image of the 3DDiscovery printer. (a) Enclosed in a flow box to print under sterile conditions. (b) A
close-up view of the 3DDiscovery: (1) flow box; (2) 3D Discovery; (3) air pressure regulators; (4) printheads;
(5) tool charger; (6) building platform; and (7) console. (c) Range of printheads and tools of the 3D Discovery.
Source : [14]
(b) (c) (a)
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why the 3D Discovery was not considered as a possibility when building the project plan. Re-
cently, RegenHU developed an experimental cartridge suitable for polymers with higher ther-
mal requirements such as PCL (Fig. 3), that fits in a DD-135N printhead. Considering the
difficulties with the bioscaffolder, it was decided to determine the potential of this cartridge for
the realization of the project.
This introduction presented the background of the project, the materials, the biofabrication ap-
proach and the tools involved. Below is the correct project process step by step:
- Implement and optimize PCL printing with the new cartridge
- Optimise the dual printing of PCL + GelMA in simple a shape in order to practice assays
- Optimise the dual printing of PCL +GelMA in an auricular shape
Therefore, for more clarity, this report is divided into 5 chapters, each of them dedicated to one
aspect of the project.
2. CHAPTER 1: PCL PRINTING
2.1 Introduction
As mentioned earlier this project is based on the ability of the printer to dual print PCL and
GelMA, thus the very first step was to set up PCL printing with the new experimental cartridge
(Fig. 3) of the 3D Discovery. The performance of this new cartridge was evaluated in relation
to the Bioscaffolder.
2.2 Material and methods.
2.2.1 Inks
Two different PCL types were used, the 704105 Polycaprolactone (Sigma Aldrich) (average
Mn 45,000) (Mw 48,000-90,000) and the medical grade PURASORB® PC 12 (Corbion Purac
Biomaterials, Spain) (IV midpoint 1.2 dl/g) (Mw 120,000[5]) (Mw = 130490, Mn = 79760
[6]).Only the 704105 Polycaprolactone from Sigma Aldrich was used with the Bioscaffolder.
2.2.2 Bioscaffolder
Manufacturing was performed at room temperature; 704105 PCL was heated up to 80°C and
printed with the extrusion printhead (Auger Screw Pump) through 330 µm (inner diameter)
nozzle (Fig.2) on the stationary platform covered with 2090 blue tape scotch 3M. PCL cannot
adhere properly to a microscope slide and deforms in absence of 2090 blue tape scotch 3M or
warming plate.
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Only the 704105 sigma PCL was printed with the bioscaffolder
since the barrel is hard to clean and the settings were already
optimized by previous work. (Precise settings in annex 1)
2.2.3 3D Discovery
In absence of the extruding printhead (HM-300H), it was neces-
sary to characterize the ability of the new cartridge to print PCL
(Fig. 3). From now on the experimental cartridge will be called
“EX”. We determined printing settings through a trial and error ap-
proach for three different nozzles; 3221130, 3221132 and
32211303 (Disposable ThinWall Insert Cores, Dispensinglink)
(Refer to: annex 1) and with two different PCL; 704105 PCL
(Sigma Aldrich), and the medical grade PURASORB PC 12 (Cor-
bion Purac). Manufacturing was performed with the EX cartridge
placed in a DD-135N printhead at room temperature, PCL heated up to 80°C and printed with
a pressure of 4.4 bar. Printing was completed on a warming plate (Thermobase platform
heater, regenHU) heating the support up to 32°C. 704105 Sigma PCL was printed on micro-
scope slides, whilst PURASORB PC 12 PCL was printed either on 2090 blue tape scotch 3M
or in a petri dish. Printing settings are present in the results. It is important to note that for both
the Bioscaffolder and the 3D Discovery the values presented in tables for “layer thickness” or
“strand interspace” are theoretical values typed into the drawing software (PrimCAM/Bio-
CAD/MMconverter). The layer thickness is the height the printer has to move up for every
additional layer, and the “strand interspace” is the distance between two strands of the same
layer but from their center (Fig. 4).
2.2.4 Characterising resolution of PCL printing - 3D Discovery versus Bioscaffolder
To characterise the resolution of PCL printing and make an accurate comparison between the
3D Discovery and the Bioscaffolder, a 2-layer design (Fig. 5 A) with decreasing strand inter-
space (1.6/1.4/1.2/1/0.8/0.6/0.4/0.2/0.1 mm) was printed, imaged with a stereomicroscope
(Olympus SZ61), and the strand thickness was calculated as the average of 10 values meas-
ured using ImageJ. In addition, PCL scaffold was printed (Fig. 5 B, C) at least 2 mm high with
a 2.25 mm strand interspace to analyze layer stacks and the overall geometry of the scaffold.
For both the 704105 sigma PCL and PURASORB PC 12 PCL, scaffolds were printed only with
the small (3221133) and medium nozzle (3221132). Pictures of the printed scaffold to measure
strand thickness can be found in Annex 2
.
Fig.4. Theoretical representation
of a PCL scaffold cross-section,
d1, d2 and d3 will be respectively
referred in our work by « meas-
ured strand thickness », « strand
interspace » and « layer thick-
ness ». Source : [8]
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2.2.5 Printing process
The printing process from modeling to printing is further explained in Chapter 4.
2.3 Results
2.3.1 Bioscaffolder - 704105 sigma PCL
The Bioscaffolder is a very reliable system for PCL printing, indeed with its extrusion-based
system, we can see (Fig. 6) that it can produce 465 µm fibers at the feed rate of 2.5 mm/sec.
Even though 2.5 mm/sec seems slow, printing PCL at such a resolution and speed is adequate.
When printing a scaffold, the bioscaffolder also performs well. We can see in Fig. 6 that the
overall square shape is true to the STL file, layers are perfectly stacked, the deposition is
homogeneous, and strands are parallel. However, in the absence of a warming plate, the print-
ing must be done on blue tape, and that is an obstacle to maintain sterility.
2.3.2 3DDiscovery
a) 704105 sigma PCL
The 704105 Sigma PCL was the first attempt to print PCL with the 3D Discovery, at that time
it was known that in absence of warming plate, the blue tape was essential to prevent PCL
scaffolds to deform during the printing. However, the team was not aware yet of the significant
impact of the support's surface on the deposition of PCL, and the thermoplate of the 3D Dis-
covery being fully functional, it was not deemed necessary to print on blue tape. In addition,
unlike blue tape, to print directly on microscope slides does not hinder sterility, and therefore,
Fig. 5. (a) two-layer design used to measure strand thickness, the strand interspace decreases on the X-axis
in the first layer and on the Y-axis in the second layer. The yellow mark indicates the first strand interspace of
1.6mm. (b) Visual representation of a PCL scaffold the red circle is pointing out what we are going to call the
single layer zone, while the green circle is pointing out the double layer zone (c) Theoretical representation of
a PCL scaffold from an upper view.
(a) (b) (c)
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was seen as an advantage. As it will be explained in the discussion, fewer results are pre-
sented for 704105 sigma PCL since efforts have been focused on PURASORB PC 12 PCL,
explaining why the printing of 704105 Sigma PCL hasn’t been further investigated on blue
tape.
We can see in Fig. 7 that 704105 sigma PCL can be printed into a range of 674,4µm to
323,5µm strands. The feed rate dropping significantly from 7 mm/sec for the biggest nozzle to
0.75 mm/sec for the smallest nozzle, changing the speed from 3x times faster than the Bi-
oscaffolder to 3x times slower. The standard deviation is reasonable for each result. If globally
the printing is satisfying; the strands are not perfectly parallel.
We can see in Fig. 7 a scaffold printed with the small nozzle 3221133, the overall shape and
geometry are satisfying, even if there are noticeable round corners. The first layer (Fig. 7) is
perfectly deposited, but since the layer thickness is not optimized, after a few layers the strand
is not perfectly stacked anymore and starts to overhang as the distance between two layers
Printer Bioscaffol-der
3DDiscovery
Material 704105 Sigma PCL
704105 Sigma PCL PURABSORB PC 12
Support Blue tape Microscope slide Petri dish Micros-cope slide
Nozzle Gauge 23 3221130 3221132 3221133 3221130 3221132 3221133 3221133
Inner Ø nozzle (µm)
337 564 335 234 564 335 234 234
Feedrate (mm/sec
2,5 ≈ 7 2 0,75 2 1 0,8 0,5 0,35 0,25
Measured strand thick-ness d1 (µm)
≈ 465,7 ≈ 674,4 ≈ 572,4 ≈ 323,5 ≈552,9 ≈242,5 ≈0,369 ≈ 186,7
≈ 280,5
≈ 437,3
Standard de-viation d1
14,74 28,68 7,5 12,65 49,37 10,06 42,2 21,73 19,63 22,72
Printer Bioscaffolder
PCL 704105 Sigma PCL
Nozzle 23 Ga
Inner Ø nozzle (µm) 337
Feedrate (mm/sec) 2,5
Strand interspace d3 (mm) 2,25
Layer thickness d2 (mm) 0,1
Number of layer 22
Construction time (min) 7,93
Length (mm) 9
Height (mm) 2,2 Fig. 6. Table of printing settings and pictures of a 704105 PCL scaffold printed with the Bioscaf-
folder from (a) a perspective view, (b) upper view and a lateral view (c)
(a) (b)
(c)
Table 1. Measurement of the dispensed strand thickness (d1) depending on the biofabrication procedure.
P a g e 15 | 78
increases, especially in the single layer zones of the construct described Fig. 5 b. Therefore,
strands start to deform and are not straight anymore, resulting in an altered shape easily visible
on a lateral view (Fig. 7c) and on the upper view (Fig. 7b). In the settings, we can read in the
table “Advised Layer Thickness (mm)”, it is a personal estimation of what should be the correct
“Layer Thickness” that would lead to an optimized construct.
b) PURASORB PC 12 PCL
All the prints of the PURASORB PC 12 have been made on a petri dish, except for one result
(Table 1). Actually, first attempts to print the PURASORB PC 12 were made on a microscope
slide with the small nozzle (3221133), exactly as we did for the 704105 sigma PCL. This ob-
tained a disappointing result since in addition to a feed rate 10 times slower than the Bioscaf-
folder, a slightly better resolution of 437 µm in average was obtained (Table 1), the strands
are not perfectly parallel (annex 1). However, when the nozzle was switched to the medium or
big size, the PCL was coming off the microscope slide, even when changing all the different
printing parameters. The conclusion was reached that the PURASORB PC 12 was not adher-
ing properly to the glass surface. After several unsuccessful attempts to increase adherence
by varying the temperature of the microscope slide, two different support were trialed, 2090
blue tape scotch 3M and a petri dish. PURASORB PC 12 shows the same printing performance
on a petri dish than on the 2090 blue tape scotch 3M, at the only difference that the PCL is
adhering so strongly on the
Printer 3D Discovery
PCL 704105 Sigma PCL
Nozzle 3221133
Inner Ø nozzle (µm) 234
Feedrate (mm/sec) 0,75
Strand interspace d3(mm) 2,25
Layer thickness d2 (mm) 0,18
Advised Layer thickness d2(mm) 0,12-0,15
Number of layer 12
Construction time (min) 14,4
Length (mm) 9
Height (mm) 2,16
Fig. 7. Table of printing settings and pictures of a 704105 PCL scaffold printed with the
3DDiscovery from (a) a perspective view, (b) upper view and a lateral view (c)
(a) (b)
(c)
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petri dish that, it can be difficult to detach it from the petri dish with no harm. Results shown
here were obtained on petri dish only because it was easier to image than on blue tape. For-
tunately, by increasing adhesion to the surface, 2090 blue tape scotch/petri dish improved the
deposition of PURASORB PC 12. We can see in Fig. 8 that PURASORB PC 12 PCL can be
printed in a range of 552.9µm to 186µm parallel strands depending on the nozzle and the feed
rate. Yet, even printed on blue tape/petri dish, feed rates used to obtain these resolutions are
significantly lower when working with PURASORB PC 12 PCL (from 2mm/sec to 0.35 mm/sec)
than with 704105 sigma PCL or the Bioscaffolder. The standard deviation is especially high for
the big nozzle (3221130) since it was hard to find optimized printing setting resulting in strands
with non-uniform width, and alteration of the overall shape. On the other hand, the standard
deviation of the sample printed with the medium nozzle (3221132) at 0.8 mm/sec is high, only
because the strand thickness is measured after imaging, and for unclear reasons the image
quality of this sample was poor, resulting in a less accurate measurement.
The printed scaffold obtained with PURASORB PC 12 PCL shows the importance of optimizing
printing settings, especially the layer thickness. For example, scaffolds printed with the small
nozzle (32211333), stacking of layers is influenced by layer thickness. In Fig. 8a, b & c the
settings are optimized, showing how well the layers are stacked, either on the single layer
zones or the double layer zones described Fig. 5b. While on Fig. 8 d, e & f, we can see
Printer 3D Discovery
PCL PURASORB 12
Nozzle 3221133
Inner Ø nozzle (µm) 234
Strand interspace d3 (mm) 2,25
Feedrate (mm/sec) ≈0,35
Support Blue tape
Layer thickness d2 (mm) 0,06 0,15
Number of layer 40 16
Construction time (min) 102,86 41,14
Length (mm) 9 9
Height (mm) 2,4 2,4
B
Fig. 8. Table of printing settings and pictures of two PURASORB PC 12 PCL scaffolds printed with
the small nozzle of the EX cartridge of the 3DDiscovery from (a)(d) a perspective view, (b)(e) upper
view and a lateral view (c)(f)
(a) (b) (e) (d)
(c) (f)
P a g e 17 | 78
overhanging between layers in the single layer zones as already described earlier for the
704105 Sigma PCL scaffold, but still correctly stacked in the double layer zones. In addition,
we can notice the impact on the overall geometry in the upper view of Fig. 9 and Fig. 9, the
first one has a perfect square shape while the second has round corners and strands are not
perfectly parallel. Furthermore, due to layers not being correctly stacked on the single layer
zones, we can reasonably suppose that the mechanical properties are strongly impeded as
well. However, it is important to point out that in addition of the printing speed being really slow,
the more layer thickness is reduced, the more layers are required to reach the same height,
and therefore the longer the printing time will be. Thus, for a simple square of 2.4 mm and
9mm of length, it required approximately 1h 43 min, so approximately 14 times longer than the
Bioscaffolder. The results concerning scaffolds printed with the medium nozzle (32211332)
don’t really provide more information and are given in annex 1. During experiments, it appeared
that in absence of warming plate, PURASORB PC 12 constructs deform on blue tape (Fig. 9)
but not on a petri dish.
2.4 Discussion
Throughout the results, we can understand how the printing output is impacted by numerous
factors, such as the support, the properties of the PCL, the nozzle or even the HMI settings. It
is hard to foresee, prior to testing, their respective degree of impact on the printing output.
Clearly shown is the importance of properly characterizing a material, especially its molecular
weight, since even though the 704105 Sigma and PURASORB PC 12 are both PCL, their
adhesion to the support and their feed rate are different. Furthermore, during research in liter-
ature we found out that in addition to changing the printing output, PCL with different molecular
weight will show different stiffness, that could have an impact on cell differentiation and poten-
tially on other factors [7]. Secondly, we observed the impact of settings such as layer thickness
on the output, from the shape of the scaffold to its mechanical integrity. Even though the over-
hanging layers described in the single layer zones could potentially be used to promote diffu-
sion at the expense of weakening the scaffold.it is important to note that if the 2090 blue tape
scotch 3M is necessary to correctly print with the PURASORB PC 12, it is an additional hurdle
for sterility in a perspective of bioprinting. Concerning the 704105 sigma PCL It is very im-
portant to take a step back, it would be a mistake to assume that the PURASORB PC 12 PCL
Fig. 9. Picture of a deformed PURASORB PC 12 PCL scaffold due to the absence of warm-
ing plate while printed on blue tape
P a g e 18 | 78
is printed with a better resolution than the 704105 Sigma PCL. It is important to keep in mind
that, by lack of time, printing settings are not optimized and the printing was studied only on a
microscope slide. Yet we know that printing on microscope slide was possible with the PURA-
SORB PC 12 only at a very slow speed, therefore, by printing the 704105 sigma PCL on blue
tape or a petri dish, theoretically it should be possible to increase the printing speed, and con-
sequently reach a resolution as good as or better than with the PURASORB PC 12.
Finally, if the PURASORB PC 12 is printed at a very slow feed rate by the 3DDiscovery, it
doesn’t necessarily mean that the 3DDiscovery is slower than the bioscaffolder, since the print-
ing of PURASORB PC 12 hasn’t been experimented with the bioscaffolder. On the other hand,
if the printing of the 704105 sigma PCL on blue tape with the 3DDiscovery is optimized, it
should be possible to increase the printing speed and consequently increase the resolution.
Furthermore, when printed on microscope slide and with the medium nozzle (32211332), that
has a comparable diameter than the bioscaffolder nozzle (335µm vs 337µm), the 704105
sigma PCL currently has a resolution only 100µm (572µm vs 465µm) higher than the bioscaf-
folder and is printed only 0.5 mm/sec slower (2mm/sec vs 2.5mm/sec) (Table 1). Therefore,
the 3DDiscovery could be as performant as the bioscaffolder.
2.5 Conclusion
In this first chapter, we saw that despite a pneumatic based extrusion the 3D Discovery is able
to properly print PCL with the EX cartridge, however, the performance thereof may highly vary
depending on the properties of the polymer itself (molecular weight) and the optimization of
printing settings and conditions.
Finally, even if this work was able to implement and optimize the printing of PCL with the 3D
Discovery and its pneumatic-based extrusion, it is important to note that the feed rate and layer
thickness necessaries to obtain good results highly vary between the different nozzles and the
PCL used. Therefore, the choice of the nozzle and of the material will condition the potential
of the 3D Discovery. With the smallest nozzle and the PURASORB PC12, the 3D Discovery
can print biocompatible PCL scaffold with precision but at the expense of a very slow printing
time, making the realization of big scaffold impossible or extremely time-consuming.
3. Chapter 2: Spraying vs Deposition
3.1 Introduction
After implementing and optimizing the printing of PCL with the 3D Discovery, the next step of
the project began: dual printing of PCL and cell-laden GelMA. However, two different
P a g e 19 | 78
approaches were possible to dispense cell-laden GelMA; spray with CF-300H (Fig. 3 C) print-
head, or deposition with DD-135N (Fig. 3 C) printhead. For better clarity, this short chapter will
be dedicated to the description of these two approaches, while the results of the experiment
are in Chapter 3.
3.1.1 Jetting of cell-laden GelMA with CF-300
Spraying liquid GelMA in between the PCL strands with the CF-300H through a microvalve is
the first method that was trialed. Firstly, the CF-300H is designed for accurate jetting or contact
dispensing, and CF stands for Cell Friendly. This method is extensively used
by several projects in the UMC laboratory, therefore it is already known as a
reliable method that can be optimized quite quickly. Since the GelMA is dis-
pensed as a liquid, optimizing the printing can be reduced to the optimization
of the quantity of material dispensed and can be done in a relatively short pe-
riod of time.
However, there are a few issues to address when printing cells with this dis-
pensing method. First, since the GelMA is liquefied at 37°C, during long printing
sessions cells slowly fall and accumulate at the bottom of the cartridge with
time (Fig. 10). As a consequence, it is really difficult to determine if the ho-
mogeneity of the cell concentration in the dispensed GelMA is preserved
throughout the printing. If gently shaking the cartridge could solve this prob-
lem, it requires unscrewing the cartridge and increase chances of breaking
sterility.
Secondly, since the GelMA is being jetted in a liquid form, it can easily be dispensed homoge-
neously in a closed volume, however it doesn’t have any shape fidelity and therefore presents
a limited scope of use as a filling, while confronting the user to issues such as leaking of GelMA
out of the printed construct.
3.1.2 Deposition of GelMA with DD-135N
Using the other approach, with the DD-135N of the 3D Discovery it is technically possible to
print GelMA in a gel form. In addition to solving issues related to the printing of liquid GelMA
mentioned earlier, it would allow new perspectives.
In the first place, cells being immobilized in the GelMA, their distribution stays homogeneous
across the cartridge and throughout the printing. Secondly and most importantly, if GelMA
could be printed with shape fidelity, it would imply the ability to precisely dispense GelMA in
complexed patterns, highly increasing the scope of use [4]. However, GelMA poor printability
Fig. 10. Picture of
the cell pellet at the
bottom of the car-
tridge in 37°C
GelMA
P a g e 20 | 78
causes difficulty for the fabrication of complex porous 3D scaffolds, and as a response, re-
searches were dedicated on its rheological properties and its dispensing. Billiet et al. showed
in their work the impact of numerous factors such as the temperature, the nozzle shape, or the
inlet pressure on the rheological properties of GelMA, and the necessity to control and monitor
these factors for precise dispensing of GelMA [8].
However, although the perspective of precisely patterning GelMA is truly interesting, it can be
really challenging to implement it, and facing these difficulties some research teams preferred
to investigate the creation of new bio inks based on GelMA that would have a better manufac-
turability [9][10].
3.1.3 Cell viability: Spraying versus Deposition
If implementing the dispensing of cell-laden GelMA in spray or deposition are two different
challenges from a technical point of view, they have a different impact too on a primary aspect
of bioprinting that is the cell viability. Preserving the cell-viability is primary in bioprinting, and
almost all the possible factors involved in the printing process were reviewed to determine how
they are impacting cell viability. Factors such as the inlet pressure, the material temperature
(directly influencing the viscosity), or the nozzle used (shape, length, inner diameter). However,
without prior knowledge, it is impossible to foresee the influence of each of these factors, and
consequently to determine which approach between spraying or deposition is likely to be the
most successful when it comes to preserving the highest cell viability. Luckily, the impact of
these factors on cell viability is well documented in the literature. Billiet et al. concluded that
the highest cell viability is obtained at a low inlet pressure (<2 bar) with conical nozzle [8]. At
high inlet pressure (>3 bar) cylindrical and conical nozzle of same inner diameter haven’t
shown any real difference. They created a heat map of the shear stress in different nozzle
(Fig. 11 B), and found that higher peak of shear stress can be found in the conical nozzle but
only at the very end of the tip, while even if shear stress is lower in the cylindrical it is present
Fig. 11. (A) Range of nozzle, (1) 300µm cylindrical nozzle for GelMA spraying, (2)(3) respectively 330µm
and 200µm cylindrical nozzle for GelMA deposition, and (4) 200µm conical nozzle for GelMA deposition.
(B) Cell-gel flow during syringe needle deposition. Heat map of the shear stress at 1 bar inlet pressure
for a conical needle (a) and cylindrical (b) needle, obtained by finite element modelling of non-cross-
linked cell-gel (10 w/v%) mixture. Fig.s are to scale for needle internal diameter of 200 µm. Source: [8]
1
2 3
4 A
A
P a g e 21 | 78
all along the nozzle [8]. This observation could explain the results obtained by Jones et al. that
concluded that cell viability diminishes with the length of a cylindrical nozzle since then cells
would be exposed for a longer time to the shear stress [11]. Finally, Yan et al. and Nair et al.
published two papers about the influence of inlet pressure and nozzle inner diameter on cell
viability, and came to the conclusion that the cell viability decreases as the pressure increases
and the nozzle diameter decreases; the effect of pressure being significantly larger than the
nozzle inner diameter. Furthermore, their surface-fitting model along with their shear stress
model shows the correlation between cell viability and shear stress induced by the process
parameters (inlet pressure and nozzle inner diameter)[12][13].
Therefore, we can try to predict the viability of our two approaches according to literature:
• In the spray approach, a 15 · 106 cells/ml cell-laden 10%(w/v) GelMA is heated at 37°C,
hence, is liquefied, and sprayed through a short cylindrical nozzle with an inner diam-
eter of 300µm (Fig. 11 A 1) and a low inlet pressure of 0.5 bar/7.25 psi.
• In the deposition approach, a 15 · 106 cells/ml cell-laden 10%(w/v) GelMA is kept at
18-24°C, hence, is viscous, and deposited through conical or cylindrical nozzle of dif-
ferent range of inner diameter (Fig. 11 A 2 3 4), with low to medium inlet pressure (1
to 3 bar).
3.2 Conclusion
Consequently, according to literature, it was expected the sprayed cells approach a higher cell
viability, since GelMA viscosity is lower, printed with a bigger nozzle and a low pressure. Even
if working with a range of different nozzles, or different temperatures, with the deposition ap-
proach, it is likely that the cell viability will be lower since the viscosity of GelMA will require
higher inlet pressure for extrusion.
4. Chapter 3: Cell viability and printing
After reviewing the main differences and features of two possible approaches for dual printing,
below outlines the printing procedure and obtained results.
4.1 Introduction
As explained in the general introduction, for dual printing a cell-laden GelMA hydrogel solution
is deposited between strands of PCL with the 3D Discovery using layer-by-layer deposition
according to a computer-aided design (Fig. 12 a). Even though the final goal of the current
work is to dual print our material in a complex ear shape, chapter 1 demonstrated that such a
print would take several hours, consequently, it would be nearly impossible to print several
P a g e 22 | 78
scaffolds in one work day. Therefore, it is difficult to optimize efficiently the deposition and is
impossible to undertake assays. Thus, as a first step, a simple design was chosen: a simple
square shape (Fig. 12 b) which consistently reduced the printing time. The GelMA will be
dispensed between strands of PCL, either as a liquid form and sprayed with a CF-300H print-
head or deposited as a gel with a DD-135N printhead (Fig. 3c). PCL is deposited with the EX
cartridge in a DD-135N printhead and with the settings used to obtain results of chapter 1. Our
first concern was to study if cell viability is preserved after these printing conditions, at which
extent and how to improve it. After facing several technical issues, a molded cell-laden GelMA
experiment that doesn’t require the correct printing of scaffolds was implemented in order to
have an insight of the impact of the printing process on cell viability despite issues related to
the material.
4.2 Material and method
4.2.1 Sterility
To assure sterile printing conditions every component in contact with printed tissue must be
sterilized, and the same guidelines used by Rimann et al. were incorporated [14]. Equipment
(thermobase, microscope slides, cartridge heater/cooler...etcetera) needed for the experiment
are installed prior to the sterilization. Special attention is given to microvalves that are cleaned
prior to sterilization by ultrasonication for 15 min at 40°C, to ensure it is not clogged. Then, the
flow box of the bioprinter was cleaned with 75% ethanol and further sterilized by UV light for
at least 30 min. Rimann et al, as well as our laboratory protocol, recommend to sterilize by
autoclave all the components to be, However, the manufacturer Nordson EFD warns on every
package to not heat the syringe barrels higher than 38°C, with no further instructions for pis-
tons, blue caps or tip caps. Thus, it was preferable to immerse every component in 75% etha-
nol and further sterilized by UV light in the flow box of the printer for at least 30 minutes.
4.2.2 Inks
Dual printing was achieved with 704105 Polycaprolactone (Sigma Aldrich) (average Mn
45,000) (Mw 48,000-90,000) and 80% DOF cell-laden GelMA. The photoinitiator combined
with GelMA in this experiment is 2-hydroxy-1-[4-(2-hydroxyethoxy) phenyl]-2-methyl-1-pro-
panone (Irgacure 2959, Ciba, Basel, Switzerland), stock solutions in PBS were filter sterilized
and mixed with GelMA to obtain a final concentration of 0.05% (w/v). The GelMA was then
dissolved in filter-sterilized PBS at the concentration of 12.5%(w/v) and mixed on a roller for 1
hour in at 37°C incubator. Confluent cells after 1 or 2 weeks of culture were detached with
trypsin to be collected, re-suspended in DMEM, centrifuged, and resuspended in PBS to be
P a g e 23 | 78
evenly mixed with GelMA (8:2, GelMA volume: cell- suspension volume) on a roller for 10 min
at 37°C, to finally obtain a final cell-laden 10% GelMA solution with a cell concentration of 15
x 106 cells/ml. ATDC5 cells were used only for training whilst equine auricular chondrocyte
cells were used for the final experiment of both printed hybrid construct and molded gels.
4.2.3 Bioprinting procedure hybrid constructs
Manufacturing was performed at room temperature, 704105 PCL was loaded and printed with
the EX cartridge placed in a DD-135N printhead and heated up to 80°C. Cell-laden (ATDC or
equine auricular chondrocyte) GelMA (see ink section) is loaded in a 10ml disposable cartridge
in a sterile manner (Nordson EFD, Ohio, USA) and co-printed either by spraying or deposition.
Printing settings can be found in Table 2. The Computer-aided design (CAD) file was designed
in BioCAD and imported in the HMI of the 3D Discovery. The two materials are co-printed on
sterile microscope slides placed on a warming plate (Thermobase platform heater, regenHU)
heated to 32°C. Once finalized, constructs were crosslinked by 10 min irradiation with a UV
Superlite (Lumatec, Munchen, Germany) (S-UV 201 A, 220 volts, 50 Hz) with an intensity of ≈
13.3 mW/cm² measured at 365 nm, a total UV irradiation of 7980 mJ/cm². Finally, constructs
were moved to 12-well culture plates containing medium and maintained in culture (see cell
culture section).
4.2.4 Cell-laden molded gel
Printing settings
PCL
Printing settings
GelMA
(sprayed)
Printing settings
GelMA
(deposited)
Material 704105 sigma PCL 10 % GelMA 10 % GelMA
Printhead 1 (DD-135N + EX) 4 (CF-300H) 2 (DD-135N)
Nozzle 3221133 cylindrical 300µm cylindrical 200µm
Pressure (bar) 4,4 0,5 3,3
Feedrate (mm/sec) 0,75 30 20
Temperature (°C) 80 37 24
Layer thickness (mm) 0,12-0,15 0,12 0,12
Strand interspace
(mm)
2,25 2,25 2,25
Cellularity (cells/ml) - 15*10^6 15*10^6
Valve opening time
(µs)
- ≈1950 -
Dosing distance (mm) - 0,7 -
Table 2. Printing settings of dual printed 704105 PCL scaffold with either sprayed GelMA or
deposited GelMA
P a g e 24 | 78
Hydrogel precursor solutions were prepared in PBS, with a final concentration of the photoin-
itiator Irgacure 2959 (Ciba, Basel, Switzerland) of 0.05% (w/v). GelMA has been dissolved on
a roller for 1 hour in a 37°C incubator. Confluent equine auricular chondrocytes after 1 or 2
weeks of culture were detached with trypsin to be collected, then re-suspended in DMEM,
centrifuged, and resuspended in PBS to be evenly mixed with 12.5 %(w/v) gelMA (8:2, gelMA
volume: cell- suspension volume) on a roller for 10 min at 37°C to reach a final concentration
of 15 · 106 cells/ml. In a sterile manner, cell-laden gelMA was either printed in deposition con-
ditions (cylindrical nozzle ø 200µm, 2.2 bar/31.9psi, 24°C) or spraying conditions (cylindrical
nozzle ø 300µm, 0.5 bar/7.25 psi, 37°C) on a petri dish, and collected to be injected in a Teflon
Mould producing gels of 6mm*2mm (disc: diameter*height) (≈56.55mm3). The gels were photo-
crosslinked using 365nm UV-light Superlite (Lumatec, Munchen, Germany) (S-UV 201 A, 220
volts, 50 Hz) with an intensity of ≈ 13.3 mW/cm² for 10 min (7980 mJ/cm²). Finally, gels were
moved from the mold to 12-well culture plates containing medium and maintained in culture
(Refer to cell culture section).
4.2.5 Cell culture and constructs culture
ATDC 5 and equine auricular chondrocyte cells were cultured in a flask maintained in a hu-
midified 5% CO2-containing atmosphere (37°C), cultivation mediums were changed every 2
days. Cultivation medium for equine auricular chondrocyte was consisting of DMEM + Gluta-
Max-I, +4,5g/L D-glucose, + pyruvate (Life Technologies), supplemented with 10% FBS, pen-
icillin (100 U/mL), streptomycin (100 µg/ mL). While ATDC5 cultivation medium was consisting
of DMEM/F-12 HAM’s medium + GlutaMax-I (Life Technologies), supplemented with 5% FBS,
penicillin (100 U/mL), streptomycin (100 µg/ mL), for ATDC 5. Confluent cells after 1 or 2
weeks of culture were collected and mixed with gelMA to be used in the biofabrication proce-
dure described earlier in this chapter. The hybrid constructs and molded gels were then placed
in 12-well plates and maintained at 37°C in 5% CO2 condition with fresh medium change every
Fig. 12. Design and three- dimensional (3D) bioprinting of hybrid constructs with structural and biolog-
ical features. (a) hybrid biofabrication using pneumatic dispenser heads. (b) Schematic of designed 3D
hybrid construct with alternating strands of polycaprolactone (PCL) and chondrocyte-impregnated
GelMA in each layer. Source [3]
(a) (b)
P a g e 25 | 78
two days. Whilst the same medium as the flask culture was used for auricular chondrocyte
constructs and moulded gel, a medium consisting of DMEM/F-12 HAM’s medium + GlutaMax-
I (Life Technologies), supplemented with 5% FBS, penicillin (10D 0 U/mL), streptomycin (100
µg/ mL), dexamethasone (0.1 mM), ascorbate-2-phosphate (0.4 mM), and ITS X (1X) was
used for ATDC 5 constructs.
4.2.6 Live/Dead assays of ATDC5 constructs and molded gels
To qualitatively determine the cell viability in the hybrid constructs and molded gels, a cell
viability assay was conducted using a LIVE/DEAD Kit (ThermoFisher scientific) and fluores-
cence microscopy. Living cells with normal intracellular esterase activity are stained by the
green fluorescent Calcein-AM dye. On the other hand, Ethidium Homodimer-1 (EthD-1) dye
enters cells with damaged membranes and undergoes an enhancement of fluorescence upon
binding to nucleic acids, thereby producing a bright red fluorescence in dead cells. The viability
of cells in the hybrid constructs and molded gels was assayed immediately after biofabrication
(day 0) and at days 1, 3, and 7 of subsequent in vitro culture. At each time point, the con-
structs/gels (n=1 per condition) were removed from the culture, washed with plain DMEM, and
stained in 2 µM calcein- AM and 2 µM EthD-1 solution in DMEM for 20 min in a 37°C, 5% CO2
incubator. The constructs were washed with PBS thrice for 5 min and imaged using an upright
fluorescence microscope (Olympus BX51) under 40X magnification. Pictures could not be
treated with ImageJ due to the high cell density and the quality of the signal, therefore quanti-
tative results couldn’t be obtained.
4.2.7 AlamarBlue® of molded gels
To quantitatively determine the cell viability in the molded gels, a cell viability assay was con-
ducted using AlamarBlue® DAL1025(Invitrogen). AlamarBlue® (resazurin) is a proven non-
toxic cell viability indicator that uses the natural reducing power of living cells to convert resaz-
urin into resorufin, therefore, the amount of fluorescence produced is proportional to the num-
ber of cells and to the metabolic activity of living cells. The viability of cells in molded gels was
assayed immediately after biofabrication (day 0) and at days 1, 3, and 7 of subsequent in vitro
culture. At each time point, the medium is changed and 1/10th volume (200µl) of AlamarBlue®
reagent is directly added to the culture media. The reaction was incubated in darkness for
16h40 at 37°C, 5%. After incubation, three 200 μl replicates were taken from each well and
transferred into a 96-well plate with a flat bottom. Finally, fluorescence is measured with a
microplate fluorometer, Fluoroskan Ascent FL (Thermo Fisher Scientific, MA, USA) (excitation
544 nm, emission 572 nm). After removing all assay solution and washing the samples twice
P a g e 26 | 78
with PBS, the samples were cultured further for use in the next experiment. The incubation
time was overnight, it is important to note that even if the assays were started at day 0-1-3-7,
results are labeled as the day after (1-2-4-8) on Fig.16 since the measures correspond to the
end of the incubation.
4.2.8 Histology staining
ATDC5 laden gelMA and PCL dual printed constructs were prepared for histological analysis,
first fixated in 4% formalin at room temperature for 24 h, then dehydrated with baths of increas-
ing concentration ethanol (70%-100%), then is subsequently be submerged in a histo-clear
bath. Finally, the samples were embedded in paraffin and cut in a 5µm slice with a microtome.
4.3 Results
4.3.1 Dual printing optimization
Unfortunately, the results are very limited as dual printing could not be optimized due to tech-
nical issues (Refer to troubleshooting). The spraying approach with 704105 PCL and sprayed
ATDC5-laden gelMA was successful and a first draft of the settings is available in table 2.
Several sprayed ATDC5 cell-laden constructs were successfully printed and were used initially
for live/dead training and optimization of the protocol. Even though the signal is not very clear
as can be seen in Fig. 13, ATDC5 can be observed on PCL strands and overall the scaffold.
Even though higher numbers of dual printed scaffolds were successfully printed, more
live/dead data is not available since for a short period the lab changed the calcein-AM (Refer
to section: troubleshooting). For the deposition approach, it is very hard to draw any conclu-
sions. Indeed, some constructs were printed with the settings provided in table 2, but the print-
ing was hardly reproducible one day to another even within the same settings and conditions,
illustrating the challenging aspect of this approach as described in the previous chapter. As-
says could not be achieved on these samples due to fungal infection. Sadly, no results of dual
printing with PURASORB PC12 PCL or with equine auricular chondrocyte are available due to
the timeline of the project and technical issue with the printer (Refer to section: troubleshoot-
ing).
Fig. 13. Live/dead picture at day 7 of dual printed construct of 704108 PCL and ATDC5 impregnated
gelMA. In red are delimited the PCL strands. Magnification x100
P a g e 27 | 78
4.3.2 Live/Dead of cell-laden molded gels
Fig. 14 demonstrates the live/dead result of the molded cell-laden gelMA pucks. At day 0, we
can observe that after printing the green signal is clear with a sharp definition of cells, there is
hardly any significant difference between each condition, it is noticeable that red and green
signals don’t overlap. The darker crack with a higher red signal on the “before printing” condi-
tion, is simply a defect in the gel due to manipulation. After day 1, in every condition the red
and green signals start to overlap, some cells presenting both red and green fluorescence.
The result of deposited gelMA at day 1 appeared to be, after further experimentation, charac-
teristic of a gel being upside down. Indeed, during incubation the gel rests on the bottom of the
Day 0
Day 1
Day 3
Day 7
Unprinted cells Sprayed cells Deposited cells
Fig. 14. Live/dead pictures of moulded gels with unprinted, sprayed or deposited chondrocyte
impregnated GelMA, from day 0 to day 7. Magnification X40.
P a g e 28 | 78
well-plate, consequently, reagents are not diffusing properly through the gel from the thereof
base. Due to a lower fluorescence signal, it was necessary at day 3 or 7 to increase the exci-
tation time, increasing at the same time the background signal coming from gelMA
autofluorescence and introducing noise into the signal. However, even if once more, it is hard
to see a significant difference between the conditions, a yellow tone can be observed but due
to the noise and the different focal planes, it is difficult to say if more red and green signal are
superposed of if there is only an increase of the red signal. We can see the same results at
the extremity of the gels Fig. 14, with a signal mostly yellow on day 7. Finally, Fig. 15 demon-
strates that in a gel cut in a half, the fluorescent signal is only present at the surface of the gel.
4.3.3 AlamarBlue® of cell-laden molded gels
Fig. 16 demonstrates the evolution of the fluorescence intensity as a function of the number
of days, the fluorescence intensity is directly related to the number of cells and their metabolic
activity. At day 1 the non-printed cells show the highest fluorescence, while the signal of
sprayed and deposited cells have a signal 3 times and 6 times lower, respectively. If we look
at numbers, sprayed cell signal is almost twice that of the deposited cells signal. Between day
1 and day 2, all the signals dropped and are at their lowest point. However, while unprinted
cells and sprayed cells signals are respectively at 62% and 50% of their initial value, deposited
cells signal dramatically dropped to 10% of its initial value. From day 4 to day 8, the signal of
unprinted cells and deposited cells are slowly rising again to reach respectively 73% and 88%
of their value on day 1. The signal of sprayed cells is more unpredictable, it suddenly increases
between day 2 and 4, to reach a value of 237% of the initial one, and finally dramatically de-
creases between day 4 and 8. Despite this second drop, the sprayed cell signal increases
again to reach a final value of 122% of the initial value (day 1). It is plausible that values meas-
ured at day 4 or 7 for sprayed cells are biased for unknown reasons, resulting in this curious
fluctuation of the fluorescence signal. Finally, on day 8 the unprinted cells signal is still the
highest with a value twice higher than the sprayed cells signal and 5 times higher than the
deposited cells. The signal of the sprayed cells was determined to be 2.5 times higher than the
deposited cells.
Fig. 15. Live/dead pictures of moulded gels with sprayed chondrocyte impregnated gelMA
at day 3 and cut in half. Magnification 40X
P a g e 29 | 78
4.4 Troubleshooting
Unfortunately, only a few results are presented here, it is important to peruse the reasons that
hindered the correct progression of the project. First, the printing of PURASORB® PC 12 is
described in the first chapter and does not appear here. It is simply because PURASORB®
PC 12 was made available in the laboratory only 2 months before the end of this internship.
Therefore, the few results presented here for the dual printing only concerns the 704105 Sigma
PCL and were obtained prior to the arrival of PURASORB PC 12 PCL in the laboratory. The
limited quantity of results with the PURASORB PC 12 PCL can be explained by the fact that
the second optimization of printing settings was necessary, retarding the implementation of
dual printing. Secondly, the cell concentration used for the project is 15.106cells/ml, however,
in order to have a reasonable amount of material (4-5 ml), it requires a substantial number of
cells and consequently a considerable culture time. This explains why the training portion of
the project was realized with ATDC5, due to their fast growth. It is essential to note that reach-
ing such a large number of cells with auricular chondrocyte can take up to 2 weeks, limiting
the number of attempts to print and obtain results at 2 or 3 print sessions only per month.
Furthermore, when the experiment was finally ready to be carried out, the microvalve was not
functional for a month. Thirdly, it is fortunate that it was possible to properly prepare the exper-
iment while cells are in culture with cell-free gelMA for the spraying approach, since at 37°C
the gelMA has a low viscosity, subsequently cells have only a small impact on the rheological
properties. Otherwise for the deposition approach where gelMA is very viscous due to a lower
manufacturing temperature, cells have a big impact on rheological properties as shown by
Billiet et al.[8], and the experiment needs to be prepared and optimized with cells. Therefore,
attempts were even more limited due to cell speed of growth for the deposition approach.
More results were also expected from the live/dead assays, unfortunately during an
experiment, the calcein regent of the laboratory was changed (aliquots labeled “Calcein-AM, 1
mg/ml, MB 21-04-2016”), this new calcein was not compatible for these live/dead assays.
When treated with the same protocol, gelMA presented an intense auto-fluorescence covering
all cells signals and made any observations impossible. The protocol was changed with differ-
ent dilutions, from ½ to 1/16 of the normal concentration of calcein, different incubation times,
trials on cell-free gels and gels with lower cell concentration, but every time the output was the
same with an intense auto-fluorescence. Eventually, a new stock of calcein from Life technol-
ogies was available in the lab allowing correctly performance of the live/dead of the last exper-
iment.
P a g e 30 | 78
As outlined in the material and methodology section, the incubation time used for AlamarBlue®
was long (16h40). Even though AlamarBlue® incubation time can go up to 18h, there might
be a more efficient incubation time. However, as printing cells is extremely time-consuming,
and AlamarBlue® incubation time takes at least 2 hours, this forced the AlamarBlue® to be
conducted overnight. The UV exposure used in the project is so high simply because the UV
intensity could be measured only at the end of the experiment when the optometer X9-2 (Gi-
gahertz-Optik, Germany) was available in the lab (Refer to annex 3).
4.5 Discussion
Henceforth, only the results obtained with the molded gels will be discussed as, with the ex-
ception of a few live/dead pictures, the dual printed samples couldn’t be used for further ex-
periments. It must be noted that due to the low number of samples and the technical issues
faced throughout this project, these results are not statistically significant, and therefore must
be interpreted with hindsight. It would be interesting to reproduce experiments a second time
with a higher sample size. The results from the live/dead assays are hard to interpret. At the
first glance, we could think that cell viability is decreasing over time and independently of the
printing process due to the yellow tone of the pictures that indicates an increase of the red
signal. However, the hypothesis of a decrease of cell viability over time would be is not corre-
lating with our AlamarBlue® results and would result in pictures with a reddish tone. If we look
closely at the pictures, we can observe that red and green signal are superposing, resulting in
this yellow tone. We know that Ethidium Homodimer-1 (EthD-1) dye produces a bright red
162,35
101,49111,18
119,31
48,49
24,62
115,18
59,64
27,83 2,9821,70 24,64
0,00
20,00
40,00
60,00
80,00
100,00
120,00
140,00
160,00
180,00
0 1 2 3 4 5 6 7 8 9
Flu
ore
sce
nce
inte
nsi
ty (
arb
itra
ry u
nit
)
Days after printing
un-printed
spraying
deposition
Fig. 16. Evolution of the fluorescent intensity of moulded gels made of chondrocyte impregnated gelMA
in unprinted, sprayed or deposited condition, as a function of time.
as function of time
P a g e 31 | 78
fluorescence upon binding to nucleic acids, but instead of signalling a damaged membrane of
a dying cell, it could indicate that cells have porous membranes. Live/Dead assay was initially
designed for two-dimensional culture, yet we know that cells can behave differently in three-
dimensional conditions. Thus, in addition of explaining the superposition of red and green sig-
nal resulting in the yellow tone of pictures, this hypothesis would be consistent with the Ala-
marBlue® results that indicate an increase of fluorescence, and consequently an increase of
cell-proliferation or cell-viability. However, the sharpness of the signal diminishing with time for
unknown reasons, and in absence of triplicate, conclusions cannot be drawn from these pic-
tures. The Live/dead assay might not be suitable for our work. The cell density is too important
and the structure too thick to allow any valuable information to be gathered in this way. Indeed,
with this number of cells, a lot of signal from different focal planes overlap and treating the
image with ImageJ is impossible. Moreover, when working with thick constructs as in this study,
it is impossible to observe all the focal planes with a fluorescence microscope. Therefore, it
would be interesting to use a confocal microscope and see if the signal is clearer when com-
piling the different focal planes. Secondly, we saw Fig. 15. that the fluorescent signal is only
present at the surface of the gel, therefore either reagent might not diffuse correctly through
the whole sample or the signal exists but can’t be observed because the light can’t diffuse
properly in the gel. Finally, the live/dead technique is a good qualitative indicator, but output
involving thick constructs with high cell density is limited.
The AlamarBlue® is a great alternative or complementary technique to the live/dead, as it gives
quantitative information on cell viability and cell proliferation despite the thickness of the
construct or a high cell density. However, it can be challenging to use it for printed scaffolds
as, to accurately compare different conditions it is important to start with an equal number of
cells between each scaffold, therefore it is of prime importance that the same volume of mate-
rial is dispensed in each scaffold and that the cells concentration is homogeneous in the car-
tridge.
In our case, the AlamarBlue® results do not support the live/dead results but matches the
literature since the spraying approach led to a higher fluorescence signal than the deposition
approach, meaning that more cells are alive or are more metabolically active. The fluorescence
signal of unprinted cells is higher than the printed cells, proving an impact of the printing pro-
cess on cell viability. All the conditions show the same tendency, the fluorescence signal of
AlamarBlue® drops in the first 48h and then recovers after the day 2. The drop of fluorescence
could be explained by the fact that apoptosis might take some time, therefore immediately after
printing cells still show an active metabolism while they are actually dying, leading to a lower
fluorescence signal 24h later. Literature has shown that if the printing process is harmless then
this drop in cell viability won’t be observed, and the cell proliferation will lead to an increase of
the fluorescence over time [4]. On the other hand, if the printing process negatively affects cell
P a g e 32 | 78
viability, we might observe a drop in cell viability and a recovery over time [16][2]. The curious
thing about these presented AlamarBlue® results is that the drop of the fluorescence signal is
present 48h after printing and then shows a recovery. Furthermore, even cells that did not go
through the printing process show a drop of the fluorescent signal. How to explain this drop for
48h and not only 12h for all the samples?
Since the drop in viability is affecting all the samples including the unprinted cells, it could be
related to the photo polymerization process that all samples underwent. The impact of a con-
centration of 0.05%(w/v) of Irgacure 2959 on cell viability is well documented in literature, and
has been shown compatible with acceptable level of cell viability in 3D construct after reason-
able UV exposure, therefore [17][18] it is unlikely that Irgacure 2959 is causing the drop of the
fluorescence signal. On the other hand, the total intensity of UV exposition in our experiment
(7980 mJ/cm²) is important, while in the literature values of UV exposition are generally around
1300-1800 mJ/cm² to preserve a good cell viability. The work of BIlliet et al. showed a cell
viability of only 55.72±7.26% for UV-A irradiation doses of 5400 mJ/cm² [8], leading to the
conclusion that an exposition of 7980 mJ/cm² must severely hinder cell viability and thus the
fluorescent signal.
Why are the fluorescence signals of the sprayed and deposited cells so low compared to the
unprinted cells signal? The answer is in the question, it may be reasonably asserted that the
printing process is the cause of the lower fluorescent signal, and even though our results are
not in triplicate they seem to concur with literature. Indeed, chapter 2 demonstrated that with
the chosen printing conditions, sprayed cells should have a higher viability than the cell depos-
ited that undergo higher shear stress during printing. However, a smaller drop of the fluores-
cent signal of the sprayed cells would have expected since the conditions and settings used
should preserve a high cell viability. The important UV exposition combined with the printing
stress could increase the adverse impact on cells. Once more, cautiousness is required re-
garding conclusions since the experiment currently has only been completed once.
Finally, in the material and method appears a histology section, even though no results are
presented in the result section. Several dual printed samples went through the described pro-
tocol, and histo-clear was used instead of xylene since xylene is a solvent of PCL and would
partly damage our samples. However, with the PCL present throughout the construct it was
impossible to properly cut a slice for staining, it could be suggested to use cryostat and see if
the slicing process is easier. On the other hand, if we would use xylene it might be possible to
slice our sample but depending on the manipulation skill of the user the sample integrity will
be affected by the empty space left by the PCL, and potentially information on the interaction
between the gelMA and the PCL would be lost. We can reasonably assume that the team of
Kang et al. that published in the magazine Nature have high-level skills, however, obtaining
P a g e 33 | 78
perfectly sliced histologic slide is very difficult when PCL is involved [15]. Therefore, the choice
between xylene and histo-clear might depend on the architecture of the PCL scaffold of the
experiment and of the skill of those involved in the project.
4.6 Conclusion and Further Experiments
The material and methods have shown the importance of correctly choosing the reagents,
even if the molecule is the same the results might be different. For instance, the calcein from
life technologies worked with gelMA constructs, however, the calcein aliquots in the lab never
gave proper results due to an intense gelMA autofluorescence that was covering any signals,
despite modification of the protocol. For any further work, it is clear that 7980 mJ/cm² UV is
harmful to cell viability and it seems logical to reduce the total UV exposition to a value closer
of 1800 mJ/cm². The next experimenter can refer to annex 3 to calculate the time of exposition
and distance from the UV-light depending on the one he will choose. Even though it is not
really necessary, potentially changing the photoinitiator would improve results, indeed Rouil-
lard et al. and Billiet et al. have both shown higher cell viability using the photoinitiator VA-086
compared to Irgacure2959[8][19].
Also discussed was the difficulty to create histology slices using the 3D printed construct partly
made of PCL, and the solvent used (xylene vs histo-clear) to prepare the histology slide must
be determined depending on experimenter’s skill and the PCL scaffold.
Secondly, related to our results, even if by lack of sample and statistical analysis our results
need to be confirmed by further experimentation, it seems to confirm that the sprayed cell
preserves a higher cell viability/proliferation than cells printed in the deposition conditions. It
was expected that sprayed cells would preserve a higher cell viability/proliferation, even though
the intense UV exposure might have biased our results. The live/dead is difficult to interpret by
lack of sharpness of the signal over time, but the superposition of the red and green signals
over time is truly interesting and could suggest a change of the membrane porosity, not related
with apoptosis. Therefore, it is recommended to reproduce the Moulded gel experimentation
with more samples to confirm the first results obtained in this study, and to use a calibration
curve for AlamarBlue® to provide better insight.
Dual printed scaffold results are discussed in Chapter 5 conclusion.
5. Chapter 4: Auricular shape and biofabrication
5.1 Introduction
The goal of this work is to dual print PCL and gelMA in the shape of auricular cartilage. How-
ever, even after going through all the previous steps we are unable to correctly dual print PCL
P a g e 34 | 78
and gelMA in a morphological shape due to software limitations. Indeed, the 3D Discovery
being a printer sold by regenHU, regenHU is in charge of providing adapted software. Never-
theless, regenHU is currently providing two software for our lab, specifically “BioCAD” and
“MMconverter”. To precisely understand this issue, it is necessary to describe these software
in more depth. Even though the Bioscaffolder is no longer a candidate for the realization of the
project, its biofabrication procedure will be described for information.
5.2 Material and method
5.2.1 Bioscaffolder and PrimCAM
The Bioscaffolder is controlled through a software combining the HMI and a drawing interface
called PrimCAM. PrimCAM provides very poor drawing tools, but its main usefulness is its
ability to import STL files and precisely control layer by layer the printing settings. On the other
hand, its interface is not perfectly adapted for dual printing and performs poorly. We can see
in Fig. 17 the typical translation of a file from its creation in a CAD software to its printing.
5.2.2 BioCAD
BioCAD is a layer by layer type of drawing software, thus it is designed to allow the creation of
simple shape with basic tools. It is possible with some expertise to create layers with a complex
internal architecture combining different material, but it is not made to design a complex shape
such as the ear. Therefore, BioCad was used in all the previous work to obtain a square-
shaped scaffold, but cannot be used for an auricular shape. We can see Fig. 18, drawings
made in BioCAD and the printed result with gelatin dyed orange.
5.2.3 MMconverter
Fig. 17. Biofabrication process of the Bioscaffolder with a STL file of the auricular shape, through
the drawing software PrimCAM, and the HMI that controls the print.
3D model (.STL) Drawing PrimCAM (.CAM)
HMI (NC code)
P a g e 35 | 78
MMconverter allows the user to import STL files and convert them into .iso files importable in
the HMI. Therefore, it is possible to work with all sort of complex shapes since the source is
an STL file that can be designed with some expertise in any 3D modeling software. The soft-
ware provides the possibility to work with different materials, but not simultaneously. “Simulta-
neously” meaning that the two materials can be printed in the same construct but not in the
same layer, which constitutes an issue. Consequently, it is possible to pattern PCL in an au-
ricular shape but impossible to precisely dispense gelMA in between the PCL strands. Further-
more, if MMconverter can import STL file like PrimCAM, it does not provide a control over the
printing settings layer-by-layer, therefore the output is considerably limited. We can see in Fig.
19 the typical translation of a file from its creation in a CAD software through to its conversion
in a .iso file understood by the 3D Discovery.
5.3 Results and troubleshooting
One issue encountered when working with BioCAD and MMconverter is the difficult translation
between the two software. BioCAD offers total control over settings such as the strand inter-
space layer-by-layer, whilst MMconverter only provides limited control options such as “filling
pattern type” and “density”. It is not clear how MMconverter calculates the strand interspace
from a filling density, and consequently, it is impossible to precisely control the strand inter-
space through MMconverter. The problem being that the printing settings controlling the quan-
tity of gelMA that needs to be dispensed mainly depends on the strand interspace, thus print
settings that have been optimized in BioCAD cannot be used directly in MMconverter until the
filling density in percentages is not directly linked to a strand interspace distance in millimeters.
For this purpose, it is necessary to choose a “density”, print a scaffold, and measure the strand
Fig. 18. Biofabrication process of the 3DDiscovery with a drawing in bioCAD to the HMI that con-
trols the print and a picture of a dual printed construct of PCL and orange dyed gelMA
Drawing BioCAD (.BCD) HMI (.iso)
Fig. 19. Biofabrication process of the 3DDiscovery with an STL file from the modelling in the CAD software
(Rhinoceros 4.0), through MMconverter, and the HMI that control the print
3D model (.STL) MM converter (.iso)
HMI (.iso)
P a g e 36 | 78
interspace manually corresponding to this filling density. An attempt was made to work around
the problem in MMconverter, since the jetting of gelMA does not necessarily require a perfect
pattern. As the jetting is not consistent, the gelMA is not homogeneous in the construct and
the PCL is not deposited correctly. This is due to the gelMA being jetted on the strands and
thus impedes adhesion on the previous layer, as the gelMA is not dyed it can be difficult to
make a sharp observation. Another foreseen problem is the printing time. Indeed, in Chapter
1 it was shown that for a simple PURASORB PC 12 PCL square of 9mm*2mm (length*height),
it took up to 1h 40 min with the smallest nozzle. However, with the platform heating the support
at 32°C, in our few attempts It has been noticed that during long prints (>1 hour) the gelMA is
drying over time. If partially crosslinked gelMA in between each layer could be a solution, it
would be much easier to simply increase the feed rate of PCL and greatly reduce the printing
time.
5.4 Discussion and Conclusion
In response to this software limitation, RegenHU developed “Object pattern”, a plugin for Bio-
CAD. It can be assumed that “Object pattern” would allow the import of an STL file while
combining the layer-by-layer precise control of BioCAD. Nonetheless, it is in beta version and
still not available in our lab. Until a new software or plugin is obtained it will be impossible to
properly dual print two materials on the same layer in a complex shape.
6. Chapter 5: General conclusion and prospective work
To end this report below is a review of the main points of each chapter in a final conclusion.
First this work demonstrated that the 3D Discovery with pneumatic based extrusion and new
EX cartridge can reliably print PCL, however, depending on the materials (material’s molecular
weight, nozzle Ø) its potential highly varies. If the 3D Discovery can print PURASORB PC 12
strand up to 186Ø it is at the expense of an extremely slow feed rate. Therefore, the 3D Dis-
covery cannot combine a high resolution and a high printing speed with its pneumatic extrusion
but remains a powerful tool. This is described as well as the importance of characterizing a
polymer, since its properties, especially its molecular weight will impact all the aspects of a
biofabrication project - from its printing to the mechanical properties of the scaffold and its
influence on cells as mentioned by Hendrikson et al.[7]Overall, the first 3D Discovery combined
with the EX cartridge remains a powerful tool despite a significantly slow feed rate as it can
combine PCL scaffolds with one material deposited with the DD-135N left and two other inks
sprayed with the two CF-300H, while the new 3DDiscovery can combine high-resolution PCL
scaffold but only with one sprayed material.
P a g e 37 | 78
This report has shown how gelMA could be dispensed in a scaffold either as liquid filling or a
shaped gel strand, and the impact of each of these approach on cell viability. The conclusion
was that the sprayed approach showed a higher preservation of cell viability after printing,
whilst being easily implementable and therefore should appear as a first choice even if as a
liquid its potential is limited. On the other hand, the deposition approach of gelMA represents
a great potential but remains highly challenging and requires the user to undertake a time-
consuming work of optimization before being efficient. This is before mentioning a bigger im-
pact on cell viability due to the higher pressure required for extrusion, and therefore should be
considered only if essential for the correct progression of a project. In the case of the gelMA
deposition, it can be interesting to consider the modification of the ink, for example as men-
tioned earlier with hyaluronic acid or gellan to increase the printability of the material.
Another aspect of this work is the importance of choosing and adapting assays depending on
the project, especially in biofabrication where depending on the sample, this may not occur.
Indeed, the materials used might not be compatible with the assay’s protocol, for example in
histology where a sample with a PCL scaffold will dissolve in xylene providing the necessity of
changing the solvent to histo-clear, which then results in samples that are hard to process with
a microtome depending on the PCL architecture of the scaffold. Beyond materials, the sample
might not be suitable for an assay due to its dimensions, this work showed a 2D sample or thin
samples where live/dead provides a lot of information. Conversely, in thick constructs, the out-
put of the assays can be greatly limited or require adaptation of the protocol conditions and
collection of the data, for example by changing the microscope. Being confronted by these
limitations can lead to rethinking the analysis of a project, in this case, for example, opting for
alamarBlue that gives quantitative results independently of the dimension of the sample and
without cytotoxicity, allowing us to experiment at different time point despite a poor number of
samples is a highly useful tool.
Finally, even when the printing process is mastered, it is important to keep in mind that the
potential of a printer directly depends on its software, in this case even if theoretically the printer
can dual print an ear shape, as the software can’t code this shape while combining two mate-
rials, it entirely hinders the progression of the project.
The conclusion of this work is that, when starting a biofabrication project, every detail of its
framework must be defined. Biofabrication provides us with powerful tools, however, the po-
tential of these tools can be fully exploited only if every aspect of the project is cautiously
studied. The numerous factors from the material itself to the software, the biofabrication pro-
cess or the assays and how they interact (Fig. 20) and interfere in the printing process can be
P a g e 38 | 78
very disconcerting for the user and lead to time-consuming and fruitless work, therefore it is
essential for the experimenter to:
- Characterize properly the materials/inks; an example provided in this report is the im-
pact of the molecular weight of two different PCL types or the physical of gelMA on the
biofabrication process
- Optimize the printing process, such as the used tools; this work showed that a different
printer or different printhead will be used depending on the resolution needed for the
PCL scaffold, the materials it must be combined with, and to the method for combining
them (sprayed or deposited). Take into account the limitations imposed by the software
that is at the origin of the biofabrication process
- Printer calen-dar
- Assay calen-dar
P a g e 39 | 78
Fig. 20. (B) Concept map of variables and relations critical to biofabrication of an hydrogel, adapted
from:[20]. (A) Simplified concept map of variables and relations critical to biofabrication of a thermo poly-
mer, (C) complex interactions between the thermo polymer and the hydrogel during dual printing, mainly
directed by the software. (D) Elements not directly impacting the biofabrication process but orbiting around
it and necessary to consider.
Thermo polymer - Polymer - Molecular weight
Printing fidelity
- Complexity
- Resolution
- Construct size
Viscosity/adhe-
rence on sup-
port
Nozzle
gauge
Fabrication
time
Software
A
C
B
Experimental assays
Planning: - Cell culture - Printer calendar - Assay calendar
D
P a g e 40 | 78
REFERENCES [1] I. a Otto, F. P. W. Melchels, X. Zhao, M. a Randolph, M. Kon, C. C. Breugem, and J. Malda, “Auricular
reconstruction using biofabrication-based tissue engineering strategies,” Biofabrication, vol. 7, no. 3, p. 032001, 2015.
[2] W. Schuurman, V. Khristov, M. W. Pot, P. R. van Weeren, W. J. a Dhert, and J. Malda, “Bioprinting of hybrid tissue constructs with tailorable mechanical properties.,” Biofabrication, vol. 3, no. 2, p. 021001, 2011.
[3] Z. Izadifar, T. Chang, W. Kulyk, X. Chen, and B. F. Eames, “Analyzing Biological Performance of 3D-Printed, Cell-Impregnated Hybrid Constructs for Cartilage Tissue Engineering,” Tissue Eng. Part C Methods, vol. 22, no. 3, p. ten.tec.2015.0307, 2016.
[4] H.-W. Kang, S. J. Lee, I. K. Ko, C. Kengla, J. J. Yoo, and A. Atala, “A 3D bioprinting system to produce human-scale tissue constructs with structural integrity,” Nat. Biotechnol., vol. 34, no. 3, pp. 312–319, 2016.
[5] L. A. Bosworth, S. R. Rathbone, R. S. Bradley, and S. H. Cartmell, “Dynamic loading of electrospun yarns guides mesenchymal stem cells towards a tendon lineage,” J. Mech. Behav. Biomed. Mater., vol. 39, pp. 175–183, 2014.
[6] E. Díaz, I. Sandonis, and M. B. Valle, “In Vitro Degradation of Poly ( caprolactone )/ nHA Composites,” vol. 2014, 2014.
[7] W. J. Hendrikson, J. Rouwkema, C. a. van Blitterswijk, and L. Moroni, “Influence of PCL molecular weight on mesenchymal stromal cell differentiation,” R. Soc. Chem. Adv., vol. 5, no. 67, pp. 54510–54516, 2015.
[8] T. Billiet, E. Gevaert, T. De Schryver, M. Cornelissen, and P. Dubruel, “The 3D printing of gelatin methacrylamide cell-laden tissue-engineered constructs with high cell viability,” Biomaterials, vol. 35, no. 1, pp. 49–62, 2014.
[9] F. P. W. Melchels, W. J. a. Dhert, D. W. Hutmacher, and J. Malda, “Development and characterisation of a new bioink for additive tissue manufacturing,” J. Mater. Chem. B, vol. 2, no. 16, p. 2282, 2014.
[10] W. Schuurman, P. a. Levett, M. W. Pot, P. R. van Weeren, W. J. a Dhert, D. W. Hutmacher, F. P. W. Melchels, T. J. Klein, and J. Malda, “Gelatin-methacrylamide hydrogels as potential biomaterials for fabrication of tissue-engineered cartilage constructs,” Macromol. Biosci., vol. 13, no. 5, pp. 551–561, 2013.
[11] A. Faulkner-Jones, C. Fyfe, D.-J. Cornelissen, J. Gardner, J. King, A. Courtney, and W. Shu, “Bioprinting of human pluripotent stem cells and their directed differentiation into hepatocyte-like cells for the generation of mini-livers in 3D.,” Biofabrication, vol. 7, no. 4, p. 044102, 2015.
[12] K. Nair, M. Gandhi, S. Khalil, K. C. Yan, M. Marcolongo, K. Barbee, and W. Sun, “Characterization of cell viability during bioprinting processes,” Biotechnol. J., vol. 4, no. 8, pp. 1168–1177, 2009.
[13] K. C. Yan, K. Paluch, K. Nair, and W. Sun, “Effects of Process Parameters on Cell Damage in a 3d Cell Printing Process,” Imece2009 Proc. Asme Int. Mech. Eng. Congr. Expo. Vol 2, pp. 75–81\n525, 2010.
[14] M. Rimann, E. Bono, H. Annaheim, M. Bleisch, and U. Graf-Hausner, “Standardized 3D Bioprinting of Soft Tissue Models with Human Primary Cells.,” J. Lab. Autom., p. 2211068214567146–, 2015.
[15] H.-W. Kang, S. J. Lee, I. K. Ko, C. Kengla, J. J. Yoo, and A. Atala, “A 3D bioprinting system to produce human-scale tissue constructs with structural integrity,” Nat. Biotechnol., vol. 34, no. 3, pp. 312–319, 2016.
[16] Y. Yu, Y. Zhang, J. a Martin, and I. T. Ozbolat, “Evaluation of cell viability and functionality in vessel-like bioprintable cell-laden tubular channels.,” J. Biomech. Eng., vol. 135, no. 9, p. 91011, 2013.
[17] N. E. Fedorovich, M. H. Oudshoorn, D. van Geemen, W. E. Hennink, J. Alblas, and W. J. a Dhert, “The effect of photopolymerization on stem cells embedded in hydrogels,” Biomaterials, vol. 30, no. 3, pp. 344–353, 2009.
[18] J. Jung and J. Oh, “Influence of photo-initiator concentration on the viability of cells encapsulated in photo-crosslinked microgels fabricated by microfluidics,” Dig. J. Nanomater. Biostructures, vol. 9, no. 2, pp. 503–509, 2014.
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[20] J. Malda, J. Visser, F. P. Melchels, T. Jüngst, W. E. Hennink, W. J. a Dhert, J. Groll, and D. W. Hutmacher, “25th anniversary article: Engineering hydrogels for biofabrication,” Adv. Mater., vol. 25, no. 36, pp. 5011–5028, 2013.
P a g e 41 | 78
ANNEX 1
Printer Bioscaffolder
PCL 704105 Sigma PCL
Needle gauge 23 Ga
Inner Ø nozzle (µm) 337
Feed rate (mm/s) 2,5
Feed XY 150
Feed Z 400
Spindle speed 200
Material Temperature (°C)
80
Support Blue tape
Printer 3D Discovery
PCL 704105 Sigma PCL
Nozzle 3221133
Inner Ø nozzle (µm) 234 Feedrate (mm/sec) 0,75
Strand interspace d3(mm) 2,25
Layer thickness d2 (mm) 0,18 Advised Layer thickness d2(mm) 0,12-0,15
Number of layers 12
Construction time (min) 14,4
Length (mm) 9
Height (mm) 2,16
Table S2. Bioscaffolder processing conditions for the deposition of 704105 Sigma
PCL.
Table S1. Technical description provided by the manufacturer of the metal nozzle used for PCL printing
with the EX cartridge of the 3DDiscovery
Fig. S1. 3DDiscovery processing conditions and pictures of scaffold printed with 704105
Sigma PCL.
P a g e 42 | 78
ANNEX 2
Fig
. S
2. S
ett
ing
s a
nd
mea
sure
d s
tra
nd th
ickne
ss o
f b
oth
th
e B
ioscaff
old
er
and
th
e 3
DD
isco
ve
ry w
ith
th
e t
wo
diffe
rent P
CL
an
d w
ith
pic
ture
of th
e 2
laye
r co
nstr
ucts
used
fo
r m
easu
rem
ents
.
P a g e 43 | 78
ANNEX 3
Table 1.
Bioscaffolder processing conditions for the 3D fiber deposition of 704105 Sigma PCL.
P a g e 44 | 78
FROM TISSUE ENGINEERING TO DRUG DELIVERY
We have seen that despite manufacturing challenges, the fabrication of complex scaffolds is
very promising for tissue engineering. However, tissue engineering is not the only field of re-
search that benefited from the development of additive manufacturing techniques together with
biomaterials compatible with those techniques. Thus, scaffolds also appeared to be potential
tools for drug delivery, and more precisely for local drug delivery. The following work explores
the usage of scaffolds as a drug carrier for local drug delivery; for this purpose, polycaprolac-
tone scaffolds were manufactured. Moreover, a novel composite was fabricated and charac-
terized to improve drug release.
P a g e 45 | 78
.
Quentin Clément PEIFFER
QUT student #: n9827102
MSc Biofabrication
Microporous polycaprolactone
scaffolds for local drug delivery
Daily supervisor: Dr.Phong Tran
Examiner: Dr.Ir.J.Malda
D/Prof.Dietmar W. HUTMACHER
Major Research Project
IHBI QUT
03/10/2016 – 03/07/2017
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LAYMAN’S SUMMARY
Drugs administered in the blood are effective but can also be very toxic as the entire body is
exposed. If delivered directly to the diseased site, drugs which are poorly soluble in the blood
can be administered in higher doses and so be more effective while being less toxic. However,
a local approach can be complicated as it implies to have access to the diseased site and as
such is especially relevant if surgery is required. As an example, local drug delivery would be
adapted for the treatment of cancer recurrence after tumor removal surgery or could be used
to reduce the risk of bacterial infections along with inflammatory reactions from the body. The
development of tissue engineering along with 3D manufacturing techniques led to the fabrica-
tion of implants with controlled features which have appeared to be promising tools for local
drug delivery. This work focus on the development of a new composite material which would
be compatible with 3D printing and that could sustain the delivery of drugs, notably by creating
small pores on the implant surface. The new composite material was successfully fabricated
in porous structures that could release drugs. Nonetheless, porous structures have shown no-
table benefits only for soluble drugs compared to the non-modified material. Finally, more work
is required to have a better understanding of how the creation of pores affects the material and
the release of the drug.
ABSTRACT
Challenges to developing efficacious systemic drug delivery systems remain, notably concern-
ing the administration of poorly soluble drugs. Local drug delivery appeared as an alternative
to avoid systemic toxicity while targeting delivery site. However, as an invasive approach, local
delivery is especially relevant in a post-operative context such as the treatment of cancer re-
currence, implant-related infections or foreign body reactions. Stimulated by tissue engineering
and the development of additive manufacturing, the fabrication of scaffolds with controlled fea-
tures have appeared to be valuable tools for local delivery. With viscoelastic properties
favorable with 3D manufacturing techniques along with a good biocompatibility, polycaprolac-
tone (PCL) is a biodegradable polymer that became prevalent in tissue engineering. The cur-
rent work aimed to characterize the printability of a novel PCL/PBS composite and investigate
its potential as a drug delivery carrier when manufactured in porous scaffolds. Rheological
measurements have shown that the addition of PBS microparticle to PCL changed its viscoe-
lastic properties in addition to modifying its sensitivity to temperature increase. Secondly, dif-
ferent release kinetics was only observed for soluble drugs while burst effect remained im-
portant despite porosity. Therefore, PCL porous scaffolds have shown limited benefits over
non-porous scaffolds, but more work is required as results suggest they could be biased by
lack of control over drug loading and drug release.
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ABBREVIATION DDS: Drug delivery system
FTIR: Fourier Transform Infrared Spectroscopy
hOB: human osteoblast
LVR: Linear viscoelastic Region
MRSA: Methicillin-resistant Staphylococcus aureus
PCL: Polycaprolactone
nPCL: non-porous PCL
pPCL: porous PCL
SSI: Surgical Site Implant
TERM: Tissue engineering and Regenerative medicine
ACKNOWLEDGMENT D/Prof. Dietmar W. Hutmacher
Dr.Phong Tran
Dr.Christina Theodoropoulos
Hoang Phuc Dang
Margaux Vigata
Tara Shabab
CARF
KEYWORDS Local drug delivery; PCL; 3D printing; Paclitaxel; Dexamethasone; Cefazolin; Vancomycin;
Porous; Scaffold; controlled release
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TABLE OF CONTENTS ABSTRACT .......................................................................................................................................... 46
ABBREVIATION ................................................................................................................................. 47
ACKNOWLEDGMENT ....................................................................................................................... 47
KEYWORDS ........................................................................................................................................ 47
1. BACKGROUND AND LITERATURE REVIEW ................................................................. 49
1.1 DDS and local drug delivery ................................................................................49
1.2 Local drug delivery of paclitaxel in adjuvant cancer treatment .............................49
1.3 Surgical site infection and implant-related infection: cefazolin and vancomycin ..50
1.4 Dexamethasone: from biocompatibility to bone regeneration ..............................51
1.5 Local drug delivery and tissue engineering: scaffolding approach .......................51
1.6 Research question ..............................................................................................52
2. MATERIAL AND METHODS ............................................................................................... 52
2.1 Material ...............................................................................................................52
2.2 Preparation of PCL/Porogen composite ..............................................................52
2.3 Rheological study ................................................................................................53
2.4 Scaffold Manufacturing .......................................................................................53
2.5 Loading efficiency and in-vitro drug release studies (Figure 2) ............................54
2.6 Normalization of drug release data......................................................................55
2.7 Fourier Transform Infrared Spectroscopy analysis of paclitaxel-loaded films ......56
2.8 Cell culture ..........................................................................................................56
2.9 Dexamethasone bioactivity studies .....................................................................56
3. RESULTS AND DISCUSSION ............................................................................................ 57
3.1 Rheology.............................................................................................................57
3.1.1 Amplitude sweep analysis: linear viscoelastic region of nPCL and pPCL .......... 57
3.1.2 Temperature sweep analysis ..................................................................................... 59
3.2 Paclitaxel FTIR and Stereomicroscopy ...............................................................61
3.3 Drug release .......................................................................................................64
3.3.1 Porous and non-porous scaffolds ............................................................................. 64
3.3.2 Drug loss and loading efficiency ................................................................................ 65
3.3.3 Paclitaxel drug release................................................................................................ 65
3.3.3 Dexamethasone: drug release .................................................................................. 68
3.3.4 Dexamethasone: bioactivity ....................................................................................... 68
3.3.5 Cefazolin drug release ................................................................................................ 72
3.3.6 Vancomycin drug release ........................................................................................... 72
4. CONCLUSION AND FUTURE WORK ............................................................................... 75
REFERENCES .................................................................................................................................... 76
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1. BACKGROUND AND LITERATURE REVIEW
1.1 DDS and local drug delivery
Drug delivery corresponds to approaches, systems, technologies, and formulations designed
to administer a pharmaceutical compound to achieve a desired therapeutic effect. Since the
introduction of the first sustained release formulation in the 1950s[1], a significant amount of
drug delivery systems (DDS) were developed and successfully translated into clinics. But de-
spite a lot of efforts, the DDS developed in the past three decades haven’t known the same
success, notably concerning their application to clinics[1] (e.g. nanoparticles)[2]. This limited
success is mainly associated with the challenges of systemic administration[3], such as the
formulations of poorly soluble drugs [1] or targeted delivery[4]. In addition to have low oral
bioavailability, poorly soluble drugs are a problem for parenteral injection as they often require
excipient which can be very toxic (e.g. Paclitaxel and Cremaphor[5]) [1] [6]. However, it is
estimated that more than 40% of new chemical entities developed in the pharmaceutical in-
dustry are practically insoluble in water [6], which make them really difficult to develop into
clinically useful formulations (e.g. paclitaxel[5]). Secondly, in absence of targeting, a drug ad-
ministered parenterally can have severe adverse effects as the drug is delivered to the entire
body (e.g. cytotoxic drug), often resulting in low therapeutic index [7] [8].
Local drug delivery was rationalized as an alternative to the systemic administration as early
as in the 1980s, notably for chemotherapy treatments [9] [10] [11]. The benefits of a local
delivery over a systemic delivery have been rationalized several times[3] [7] [12]. By being
delivered directly to the targeted site, drugs can be administered at higher doses, and so, be
more potent while reducing the whole body toxicity [12]. Besides increasing the therapeutic
index, local drug delivery systems are also associated with longer exposure time as the drug
release is sustained for a prolonged period of time[3]. Nonetheless, except for parts of the body
which are easily accessible from the outside, e.g. skin, local delivery implies an invasive inter-
vention. However, the development of tissue engineering and the increasing use of medical
devices tend to indicate that local delivery is very promising. Below are presented 3 clinical
issues that could be addressed with local delivery.
1.2 Local drug delivery of paclitaxel in adjuvant cancer treatment
Paclitaxel is a chemotherapy drug from the Taxane family that stabilizes the microtubule poly-
mer and protects it from disassembly, ultimately leading to cell apoptosis during mitotic cellular
function [13]. Paclitaxel has a proven potency and is used in clinics to treat various cancers
such as ovarian cancer [14], metastatic breast cancer[15], or anthracycline-resistant breast
cancer[16]. Yet, paclitaxel is poorly soluble and its systemic administration is problematic and
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is associated with a low therapeutic index and an important toxicity[17]. The need for new
formulations of paclitaxel has been rapidly acknowledged [1][18][19], and similarly to other
toxic chemotherapy agents, the localized delivery of paclitaxel was suggested as a possible
solution [5] [20]. Local delivery would be extremely relevant in the adjuvant treatment of breast
cancer since an implant could be placed during surgery. Paclitaxel has already shown some
efficacy in the adjuvant treatment of breast cancer to prevent local recurrence which is a major
cause of death [21] [22]. An implant could deliver a high dose of paclitaxel increasing its bioa-
vailability to the tumor site while minimizing systemic side effects and improve the long-term
survival and quality of life of patients [7]. Furthermore, this implant could also deliver other
therapeutic agents to prevent bacterial infection or improve healing as explained below, nota-
bly when the risk of infection could possibly influence cancer recurrence[23].
1.3 Surgical site infection and implant-related infection: cefazolin and vancomy-
cin
With an incidence of 38%, SSI is the most common nosocomial infection among the surgical
patient, while accounting for approximately 15%[24] (up to 20% depending on the study[25])
among hospitalized patients, making them the third most frequently reported nosocomial in-
fections[25]. Thus, SSI and its complications represent an important clinical and economic
burden[24][25]. The risk of SSI can persist for up to 30 days after a surgical operation or even
up to one year if the patient is given an implant. Thus, the use of an implant and medical device
suffer from additional risks of ‘‘device-related’’ or ‘‘implant-associated’’ infection[3]. During
surgery, pathogens can colonize an implant surface which can result in host patient morbidity,
device removal and mortality in some cases. It appeared that preventing bacterial adhesion to
an implant in the next few hours following surgery, notably to avoid the formation of biofilms,
is decisive to achieve a long-term success of an implant [26] [27]. Hence, it was postulated
that the presence or release of antibiotics from the implant surface could possibly prevent such
colonization and significantly reduce the risk of implant-related infection[3].Cefazolin and van-
comycin are antibiotics commonly used in clinics as perioperative prophylaxis treatment to
prevent surgical site infection. Cefazolin is known for its efficacy against gram-positive bacteria
such as Staphylococcus aureus responsible for a large amount of SSI [25][28][29], while van-
comycin is meant to treat or prevent severe infection and is notably used against Methicillin-
resistant Staphylococcus aureus (MRSA) infections.[29] [30]. As such, the local delivery of
vancomycin and cefazolin to prevent SSI and implant-related infection, and more precisely
their efficacy to inhibit biofilm formation were studied and have shown promising results [3] [26]
[31] [32]. Venkata P. Mantripragada et al. even combined cefazolin and vancomycin to be
released together from microparticle for a better antibacterial effect [33].
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1.4 Dexamethasone: from biocompatibility to bone regeneration
When it comes to the clinical applications of biomaterials and medical devices, one major chal-
lenge remains beyond the risk of infection; the inflammatory reaction of the host, and notably
the foreign body reaction [34]. Any implant can trigger a non-specific host response leading to
the formation of a fibrous capsule surrounding the implanted object, which can lead to failure
of the implant [35]. As a result lot of efforts were done to have a better understanding of the
mechanism regulating biomaterial biocompatibility [36]. To improve the overall host response
and implant integration it has been proposed to reduce the inflammatory reaction. Dexame-
thasone is a corticosteroid medication used in diverse conditions to treat inflammation and
have shown promising results for such application. Hence, incorporation of controlled-release
systems, notably PLGA microsphere, of anti-inflammatory drugs such as dexamethasone was
developed and are very promising [3] [37] [38] [39] [40]. But local delivery of dexamethasone
could also potentially serve another purpose in TERM. In addition to anti-inflammatory proper-
ties, dexamethasone is known to be used in cell culture for its potent osteogenic differentiation
effect [41] [42]. Therefore, it has been suggested to combine bone implant with the controlled-
release system to deliver dexamethasone, and promising results seem to indicate that dexa-
methasone can improve overall bone regeneration [43].
1.5 Local drug delivery and tissue engineering: scaffolding approach
If local delivery from an implant is promising, it is a challenge to get precise control over the
kinetics of the release. The local release will depend on the implant and drug properties and
how they are combined, in this way tissue engineering highly stimulated drug delivery research
[44]. The development of biodegradable polymers [45] [46] [47] together with the emergence
of 3D manufacturing techniques led to the fabrication of scaffolds as novel therapeutic tools
for local drug delivery [48][49][50]. Polycaprolactone (PCL) is one of the biodegradable
synthetic polymers which has been increasingly used in tissue engineering and drug delivery
research. In addition, to be biocompatible, its viscoelastic properties and low melting point are
compatible with various manufacturing techniques which makes it an ideal polymer[51][52].
The combination of PCL with different manufacturing techniques to fabricate new PCL-based
delivery systems of therapeutic agents (drugs, DNA, and siRNA, proteins, growth factors) have
been extensively studied [53];[54], and summarised by Debasish Mondal et al.[55]. Several
strategies have been developed to load therapeutic agents on such DDS. On one hand,
therapeutic agents are encapsulated and/or blended in the matrix of the scaffold[55][56], the
release kinetics is then depending on the degradation rate of the polymer[57]. This approach
has been shown to be promising and is notably used with PCL to achieve a prolonged release
and avoid a burst effect by using PCL slow biodegradability. Still, this approach can require
P a g e 52 | 78
sophisticated manufacturing techniques as it can potentially highly decrease drug bioactivity
by exposing it to the manufacturing process (e.g. high temperature, organic solvent etc..)[56].
On the other hand, therapeutic agents can be deposited on the surface of scaffolds as a
coating, in this case, the release kinetic is depending on the degradation of the scaffold
together with a diffusion process[57]. This approach is favorable to the release of a higher dose
of therapeutic agents in a shorter period of time. The deposition of therapeutic agents on the
scaffolds can be seen as more versatile and can preserve higher rate of bioactivity for unstable
drugs as the process is independent of the manufacturing of the scaffold. Yet, this approach is
also known to be associated with an important burst release of therapeutic agents and a poor
ability to sustain the release over a short period of time [57]. Previous studies have shown that
manufacturing scaffolds with micropores was a promising technique[56]. By increasing the
surface area, a higher amount of therapeutic agents can be loaded and released in more
sustainable ways, it also seems that porous surfaces also to favor initial cell adhesion by
stimulating cell anchorage [56].
1.6 Research question
This work aimed to investigate the potential of porous PCL scaffolds as drug carrier compared
to non-porous scaffolds. To this end, a novel PCL/Porogen composite was fabricated and char-
acterized with rheology to determine its viscoelastic properties and printability. The composite
was then manufactured with 3D printing technique into porous scaffolds with an increased
surface area with the aim to improve drug loading efficacy and reduce burst release. Finally,
the drug release kinetics of 4 different therapeutics agents was studied; Paclitaxel, Dexame-
thasone, Cefazolin, and Vancomycin.
2. MATERIAL AND METHODS
2.1 Material
- Medical grade polycaprolactone Purac Purasorb PC12 (Mw 120,000 g/mol)
- PBS (Oxoid, BR0014 Dulbecco `A’ Tablets)
- Paclitaxel powder from Taxus brevifolia, ≥95% (HPLC) (T7402-5MG, Sigma-Aldrich)
- Dexamethasone powder, ≥98% (HPLC) (D1756-100MG, Sigma-Aldrich)
- Vancomycin hydrochloride from Streptomyces orientalis (V2002-1G, Sigma-Aldrich)
- Cefazolin sodium salt 89.1-110.1% (C5020-100MG, Sigma-Aldrich)
2.2 Preparation of PCL/Porogen composite
To prepare the composites, PBS pellets were crushed in a powder with a pestle and mortar.
The PBS powder was then sieved to obtain microparticles with Ø≤ 38µm. Three blends of PCL
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and porogen of different ratios (17/33/44 w%) are dissolved together in chloroform to be sol-
vent-casted in a beaker. After cooling down, the films are peeled and placed in zip bags to-
gether with autoclave bag containing silica gel beads. The bags are left open in the fume hood
for 2 days after what they are stored in a low humidity chamber.
2.3 Rheological study
It is important to precise that rheological measurements were done on PCL/PBS 17/33/44 w%
composites films which are not leached, as our focus is on how PBS microparticles are influ-
encing the printability of PCL. Rheological measurements were carried out on a rheometer
(Anton Parr M302 Rheometer) equipped with a cylindrical, parallel plate geometry using 25
mm diameter plate and 1mm gap. The different composite films were melted at 110°C for 5
min and underwent amplitude sweep tests to investigate their linear viscoelastic region. The
shear strain was ranging from 0.01% to 150%, and an angular frequency of 10 rad/sec was
kept constant during the analysis. Temperature sweep tests were also carried out on the dif-
ferent composite films to have a better understanding of their viscoelastic behavior in function
of the temperature. Composite films were melted at 125°C and the temperature was gradually
decreased by 5°C between each measurement with a 5min delay until reaching the final tem-
perature of 40°C. The angular frequency was maintained to 10 rad/sec and the shear strain at
1% during the analysis. For both temperature and amplitude sweeps, G' and G" were plotted
in two different figures for a better visibility, but to be able to compare them properly, a third
figure plotting tan δ was realized. Tan δ (also called loss tangent) corresponds to the ratio of
loss modulus to the storage modulus.
𝑇𝑎𝑛 𝛿 = 𝐺′′
𝐺′
2.4 Scaffold Manufacturing
Manufacturing was performed at room temperature. Medical grade PC 12 PCL (nPCL) and the
PCL/PBS composites films were respectively heated up to 100°C and 110°C for 30 min and
printed with a screw-driven extrusion-based printer (Fig. 1 a, b) through a ≈413μm ID (inner
diameter) nozzle on a stationary platform. Printed scaffolds measure 40mm2 and possess 6
layers, each layer being printed orthogonally to the previous one. The scaffolds were then
leached for 14 days in 0.01M NaOH to remove the PBS microparticles, creating different po-
rosities depending on the porogen ratio that was used. We can see (Fig. 1 c, d) a pPCL 44w%
scaffold after leaching. For all the experiments involving scaffolds, only scaffolds made of nPCL
and pPCL 44w% will be used. For more clarity, leached porous scaffolds and films will be
P a g e 54 | 78
designated as pPCL 17/33/44 w% for the different composites, while nPCL will stand for non-
porous PCL.
2.5 Loading efficiency and in-vitro drug release studies (Figure 2)
To carry out drug release experiments, scaffolds are cut with a scalpel in small scaffolds of
approximately 3 mm x 3 mm x 3mm (Fig.1 c, d). Those scaffolds are then weighted and placed
in 2ml microcentrifuge tube. The microcentrifuge tubes containing the scaffolds are sterilized
with 100% ethanol and dried for 24h in a flow cabinet. Release studies were carried out for 4
different drugs: Paclitaxel, Dexamethasone, Cefazolin, and Vancomycin. The quantity of drug
pipetted for loading is normalized by the weight of each scaffold. Therefore, Paclitaxel,
Cefazolin, and Vancomycin are loaded with three different doses; low dose: 0.4µg/mg of
scaffold, medium dose: 2µg/mg of scaffold, high dose: 10µg/mg of scaffold. On the other hand,
dexamethasone is loaded with three different doses: low dose: 1µg/mg of scaffold, medium
dose: 5µg/mg of scaffold, high dose: 25µg/mg of scaffold. Paclitaxel and dexamethasone were
dissolved in 100% ethanol while cefazolin was dissolved in 90% ethanol, and vancomycin in
70% ethanol, drugs were then pipetted in the 2ml microcentrifuge tubes containing the scaf-
folds and dried for 48h in a flow cabinet. Before carrying on the release studies, the scaffolds
were transferred to new sterile microcentrifuge tubes. The first tubes were stored at 4°C and
are used to measure the drug lost during loading. A summary of the protocol is schematized
Fig 1. (a) Schematic representation of screw-driven extrusion-based print head reprinted from
Valkenaers et al. [65]. (b) Screw driven Extrusion-based printer of IHBI biofabrication facility used
to fabricate PCL scaffolds. (c) (d) Top and lateral view of a PCL/PBS 44w% scaffold after leaching,
red bars correspond to 1 mm.
(b)
(c) (d)
(a)
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Fig. 2 To measure dexamethasone and paclitaxel drug loss, 100% ethanol is added to micro-
centrifuge tubes, then vortexed and sonicated for 5min, ethanol is then diluted to 50% with
PBS to avoid a rapid evaporation. For Vancomycin and cefazolin, PBS is directly added to the
microcentrifuge tubes which are then vortexed and sonicated for 5min. Solutions are then col-
lected in UV-vis 96-well plate and measured with spectrophotometry (Bio-Rad, xMark). Dual
wavelength was used to measure drug concentration: Paclitaxel (230-228 nm), Vancomycin
(282-232 nm), Cefazolin (272-254 nm), Dexamethasone (235-241 nm).
Concerning the drug release, release solutions are pipetted in the tubes containing the scaf-
folds and placed in an orbital shaker incubator at 37°C (New Brunswick™ I26, Eppendorf).
Paclitaxel is released in 200µl of a PBS/Tween 20 0.1% mix due to its very poor solubility.
Cefazolin and Vancomycin are released in 200µl of PBS, and Dexamethasone is release in
2ml of PBS to avoid sink condition. The drug release is measured in UV-vis 96-well plates with
spectrophotometry similarly to the drug loss.
2.6 Normalization of drug release data
Three experimental repeats were done for each drug release study. Despite trying to be as
consistent as possible during each repeat, it appeared that the absolute quantity of drug re-
leased between experimental repeats can importantly vary depending on the drug. Those dif-
ferences between experimental repeats could be due to small variations of drug concentration
while preparing the stock solution for drug loading. But it is also possible that, for unknown
Fig 2. (a) Schematic representation of the loading process, (b) the drug loss measurement, (c)
drug release experiment and (d) the drug-scaffold bond characterization.
P a g e 56 | 78
reasons, the drug loading efficiency vary between experimental repeats. Hence, in absence of
precise control over the amount of drug released by the scaffolds no conclusions can be made
from the absolute value. Consequently, it was decided to focus on the release kinetics of each
drug rather than absolute values. To do so, data have been normalized by the overall quantity
of drug released.
2.7 Fourier Transform Infrared Spectroscopy (FTIR) analysis of paclitaxel-
loaded films
An FT-IR Spectroscopy analysis was performed to have a better understanding of the interac-
tion between paclitaxel and PCL. The FTIR analysis was carried out on both scaffolds and
films by nPCL and pPCL 44% with a Nicolet iS50 ATR-FTIR. Both scaffolds and films were
leached for 14 days in 0.01M NaOH before being loaded with paclitaxel, similarly to the drug
release experiments. The FTIR analysis was performed over the region 4000-400cm-1, each
spectrum is the average of 64 scans; results are presented in Fig. 6. The nPCL and pPCL
films used for the FTIR analysis were also observed with a stereomicroscope. Table 1 is based
on literature [58] [59] [60] and summarizes all the vibrational bands which could be potentially
associated with paclitaxel.
2.8 Cell culture
A primary cell line of human osteoblast (hOB) isolated from human's tibia bone were cultured
in flasks maintained in a humidified 5% CO2-containing atmosphere (37°C). The culture me-
dium was changed every 2 to 3 days and consisted of MEM-α with nucleotides and nucleosides
(Gibco, catalog number 12571), supplemented with 10% FBS, penicillin (100 U/mL), strepto-
mycin (100 μg/mL).
2.9 Dexamethasone bioactivity studies
Two cell experiments were carried out to determine the bioactivity of dexamethasone. The first
experiment was performed with free dexamethasone, while the second experiment used dex-
amethasone loaded scaffolds. In both experiment, 10^3 cells per well were seeded in 96 well
plate. In the first experiment, 24h after the cells were seeded, different doses of free dexame-
thasone were directly added to the culture media (from 260µg to 0.25µg). After 24 hours of
treatment, cell viability was assessed by Alamar blue and light microscopy pictures were taken.
In the second experiment, cell viability was assessed when exposed to pPCL 44w% and nPCL
scaffolds loaded with three doses of dexamethasone (150µg/scaffold; 30µg/scaffold; 6µg/scaf-
fold). Unlike drug release studies, the quantity of drug loaded is not normalized by the weight
of the scaffold, as it is important to have the same concentration of drug between samples.
P a g e 57 | 78
Cell viability was then measured by Alamar blue at D1, D3, and D5. Values measured with
Alamar blue were reported to an approximate cell number by using a standard curve.
3. RESULTS AND DISCUSSION
3.1 Rheology
Rheology is a powerful tool to study the viscoelastic behavior of a polymer. In this study, the
rheological analysis was carried out on the different composites to determine how the PBS
microparticles content affects the viscoelastic behavior of PCL. Bear in mind that G’’ and G’
respectively describes the viscous and elastic properties of a material. Tan (δ) corresponds
to the ratio of G’’ to G’, and as a result, depending on if tan (δ) tends to 0 or infinity, the closer
the samples resemble the properties of a pure solid or a pure fluid, respectively. In Fig. 3 & 4,
G’’ always has superior values than G’. This can be explained by the fact that measurements
are carried out at temperatures above the melting point of our polymers. Consequently, sam-
ples are melted and have viscous-dominant (liquid-like) properties, rather than elastic-domi-
nant (solid-like) properties, resulting in G’’ being superior to G’ and Tan (δ) being ≥1.
3.1.1 Amplitude sweep analysis: linear viscoelastic region of nPCL and pPCL
Amplitude sweep measurements were conducted to determine how the PBS microparticles
influence the linear viscoelastic region of neat PCL. Fig. 3 (a) & (b) respectively show the
storage modulus (G') and loss modulus (G") in function of the strain. Similarly to literature a
linear viscoelastic region (LVR) can be observed for neat PCL[61]. While this LVR is relatively
conserved for pPCL17%, this is not the case for PCL/Porogen 33% and 44%. The more the
content of PBS microparticles increases the less PCL shows linear viscoelastic properties. We
can also observe Fig. 3 & 4 that G’’ and G’ values increase with the amount of PCL micropar-
ticles. This trend has also already been observed in previous works with PCL/Silica nanocom-
posites [61]. However, G’ and G’’ are influenced differently depending on the amount of PBS
microparticles and the shear strain. Hence, it is more accurate to interpret the variations of G’’
and G’ together by plotting Tan (δ). First, we can see Fig. 3 (c) that tan (δ) is linear for neat
PCL, which means that G’’ and G’ stay relatively constant independently of the shear strain.
This linear behavior suggests that the internal friction of neat PCL is independent of the strain.
On the other hand, Tan (δ) loses its linear nature with higher content of porogen microparticles,
indicating that G’’ and G’ are influenced differently by the shear strain. Tan (δ) gets closer to 1
at low shear strain, meaning that G’ knows a bigger increase of its value relatively to G’’. Hence,
at low strain and at the same temperature, composites with higher PBS microparticles are
featuring more solid-like properties compared to neat PCL. And the fact that Tan (δ) increases
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Fig 3. Amplitude sweep tests results. G’, G’’ and Tan (δ) are respectively plotted in (a), (b) and (c) in function
of the shear strain (%) from 0.01% to 150%. Temperature was kept constant at 110°C and angular frequency
at 10 rad/sec. Each point is the average of three experimental repeats (N=3); the bars correspond to SD.
(a)
(c)
(b)
P a g e 59 | 78
with the shear strain, suggests that the properties of PCL/PBS composites get more and more
liquid-like as the strain increases until they finally reach the same state than neat PCL as the
structure starts to break down.
3.1.2 Temperature sweep analysis
In Fig. 4, G", G' and tan(δ) are plotted against the temperature from 125°C to 40°C, while shear
strain is kept constant at 1%. This analysis was carried out to determine how the PBS micro-
particles content influences G", G' or Tan (δ) depending on the temperature. When looking at
temperature sweep results, the first noticeable thing is that G’ and G’’ are linearly decreasing
with higher temperatures. These results can be rationalized by the fact that, as the temperature
gets higher the polymer get progressively closer and closer to a liquid-like state. Hence, vis-
cosity (G’’) and elasticity (G’) diminish as less force is required for the deformation of the ma-
terial. Similarly, to amplitude sweep results, G’ and G’’ have higher values as the amount of
PBS microparticles increases, yet again plotting Tan (δ) is necessary for further interpretation.
Thus, for a constant strain of 1% and between 55°C to 120°C Tan (δ) is lower for composites
with higher PBS microparticles content, which comparably to amplitude sweep results suggest
that those composites feature more solid-like properties. Interestingly, tan (δ) of PCL/PBS 33%
& 44% is linear and less steep than for neat PCL or PCL/PBS 17%, which on the other hand,
seem to be exponentials. These results suggest that neat PCL and PCL/PBS 17% viscoelastic
behaviors get exponentially close to a liquid as temperature rise. On the other hand, higher
PBS microparticles contents seem to make PCL less sensitive to temperature changes; and
at 110°C Tan (δ) of PCL/PBS 44% is almost twice lower than neat PCL. This result is compat-
ible with the fact that higher temperature is needed to print composites with higher microparti-
cles content. Finally, at low temperatures, a G' and G" crossover can be observed Fig. 4. (c).
This crossover means the material starts to behave like a viscoelastic solid and not like a melt
anymore. For this reason, measurements below 50°C are erratic because the material fully
cooled down and that the rheometer is unable to take a proper measurement. Interestingly,
the G'/G" crossover takes place between 55-50°C for the PCL/PBS 33% and 44%, while it
takes place between 50-45°C for neat PCL and PCL/PBS 17%. The temperature sweep anal-
ysis is achieved with decreasing temperatures as mentioned in material and methods. Thus, a
crossover point of G’/G’’ at a higher temperature would suggest that PCL cools down faster
with a high content of PBS microparticles. This behavior could be rationalized by the fact that
PBS microparticles reduce PCL particles freedom of movement which consequently cools
down faster.
P a g e 60 | 78
Fig 4. Temperature sweep results. G’, G’’ and Tan (δ) are respectively plotted in (a), (b) and (c) in
function of the temperature from 125°C to 35°C. Strain was kept constant at 1% and angular frequency
at 10 rad/sec. Each point is the average of three experimental repeats (N=3); the bars correspond to SD.
(c)
(b)
(a)
P a g e 61 | 78
3.2 Paclitaxel FTIR and Stereomicroscopy
The infrared spectrum of a compound is formed by the superposition of the absorption bands
of its specific functional groups and as such can be used as a fingerprint. Measurements were
first taken on porous paclitaxel-loaded scaffolds to be as close as possible of the experimental
conditions of the drug release experiment. However, results weren’t reproducible, which could
be linked to the architecture of scaffolds reducing the contact surface with the FTIR probe.
When using films rather than scaffolds, no paclitaxel signal could be detected on porous scaf-
folds either, even though the deposition of paclitaxel was confirmed by a drug loss measure-
ment (data not shown). We can see on stereomicroscope results Fig. 5 (c) & (d) that no dif-
ferences can be observed between the paclitaxel-loaded and the non-loaded porous film. On
the other hand, the deposition of paclitaxel is easily observable on non-porous film Fig. 5 (c)
& (d). This absence of a signal on porous films could be due to the paclitaxel being deposited
more homogeneously in the pores of the scaffold, resulting in a low drug signal on the surface
of the film. Paclitaxel signal was detectable on non-porous films only when the FTIR probe was
placed on the white drug clusters Fig. 5 (b). Consequently, the results presented are coming
from measurements taken on those white clusters of paclitaxel present on non-porous films.
The spectra presented in Fig 6. are discussed as follow. The spectra of PCL Fig. 6 (a) is
similar to literature [62] [63]. Thus, we can observe the peaks commonly identified such as the
methyl/methylene C-H saturated aliphatic groups asymmetric and symmetric stretching re-
spectively found at 2943 cm-1 and at 2864 cm-1. The strong absorption peak at 1720 cm-1 is
Fig 5. Stereomicroscopy pictures of non-porous (a) (b) films and porous (c) (d) films before (a) (c) and
after (b) (d) loading with paclitaxel. The red bars correspond to 1 mm.
(c)
(a) (b)
(d)
P a g e 62 | 78
assigned to the stretching vibration of the carbonyl compounds -C=O, and the signal at
1470.92 cm-1 seems to correspond to the bend of the methyl/methylene C-H saturated ali-
phatic. Finally, the signal at 1162 cm-1 and 730 cm-1 can be respectively assigned to the alkyl-
substituted ether -C-O-C- stretching and the rocking motion of the methylene. Fig. 6 (e) corre-
spond to the Paclitaxel spectra and is also comparable to what can be found in the literature
[59] [60]. Nonetheless, some bands can be attributed to different functional groups and differ-
ent interpretations were given. The wide vibrational band spreading from 3600cm-1 to 3200cm-
1 correspond to different functional groups such as the O-H hydroxyl stretching and the N-H
amine stretching. Some peaks in the 3130 cm-1 – 3070 cm-1 range could correspond to the C-
H aromatic stretching. Comparably to PCL, the two peaks at 2962 cm-1 – 2888 cm-1 are re-
spectively associated with the saturated aliphatic CH asymmetric and symmetric stretching
vibrations, while the peaks found at 1730 cm-1 and 1702 cm-1 are associated to the -C=O car-
bonyl groups (ester, ketone). The band at 1635 cm-1 is assigned to the amide C=O-N bound.
Literature has previously assigned the 1602 cm-1 band specifically to the aromatic C=C stretch-
ing vibration [60], but the next peaks at 1578 cm-1 and 1533 cm-1 are harder to identify and can
correspond to the amine N-H bends but also to the C=C-C aromatic stretching. The band at
1491 cm-1 is usually assigned to the bend of the methyl/methylene C-H saturated aliphatic
bend like for PCL. The signal at 1273 cm-1 has been previously attributed to the amine C-N
stretching [60] [59] and the strong band at 1241 cm-1 to the C-O ester bond stretching vibration
[60], which could be compatible with the peak found at 1251 cm-1 in the current work. Several
Table 1. Vibrational band assignments of Paclitaxel drug
P a g e 63 | 78
Figure 6. FTIR spectra of PCL (a) Paclitaxel (e) and non-porous PCL films loaded
with low dose (b) medium dose (c) and high dose (d) of paclitaxel
(a)
(b)
(c)
(d)
(e)
P a g e 64 | 78
functional groups can correspond to the band found at 1072 cm-1 and different interpretations
were given in the literature. Thus, it was assigned to the secondary alcohol C-O stretching [59]
or to the C-H out of plane deformation bands [60]. But the peak might be more likely to corre-
sponds to the secondary alcohol C-O stretching which can be assigned to a wide range of
absorption around 1150 cm-1 and 1000 cm-1 [58], while the C-H out of plane deformation
bands are assigned to the range of 900 – 670 cm-1 in general [58]. Nonetheless, it could cor-
respond to the C-H aromatic in-plane bend which is found in the range of 1225 cm1 to 950 cm-
1[58], and that can be assigned to several peaks, such as the band at 981 cm-1. Finally, the
band at 707 cm-1 and the surroundings ones are believed to correspond to the C-H aromatic
out of plane bend [58] [60], but could also possibly correspond to the O-H alcohol out-of-plane
bend[58] The peaks 707 cm-1/981 cm-1/1635 cm-1/1578 cm-1/1533 cm-1 will be specifically
kept as fingerprint of paclitaxel for our work. Fig 6. (b), (c) & (d) correspond to non-porous
PCL films respectively loaded with a low, medium, and a high dose of paclitaxel. The first thing
to notice is that no paclitaxel signal is detected for the low dose as the spectra is identical to
the one of PCL. On the other hand, we can clearly see that the bands associated with paclitaxel
and circled in red are becoming sharper as the dose increase. However, the bands visible in
Fig 6. (b), (c) & (d) are at the same positions than on the spectra of PCL and paclitaxel ob-
served separately. This absence of a shift is generally associated with an absence of
interaction between the two materials as no functional groups bonded between paclitaxel and
PCL.
3.3 Drug release
3.3.1 Porous and non-porous scaffolds
As mentioned in material and methods, all the scaffolds used for drug release experiments are
first leached for 14 days and then cut by hand with a scalpel to obtain small scaffolds of ap-
proximately 3mmx3mmx3mm. The quantity of drug loaded in each release study is normalized
Figure 7. Bar diagram representing the difference of weight between porous and non-
porous scaffolds of similar size after leaching. The bars correspond to SD (n=146)
P a g e 65 | 78
by the weight of each scaffold. For this reason, each scaffold is weight before the sterilization
and loading process. If more specific techniques are required to study the porosity of porous
scaffolds in details, we can already see Fig. 7 the significant difference in weight between
porous and non-porous scaffolds of the same size. Thus, porous scaffolds and non-porous
scaffolds respectively weight 5.6 mg and 9mg in average. Furthermore, it has been observed
during experimentation that porous scaffolds are floating in PBS
3.3.2 Drug loss and loading efficiency
The drug loss and loading efficiency were measured and calculated for all drug release exper-
iments as explained in material and methods. Unfortunately, the results can’t be shown as they
weren’t reproducible. It is not clear yet if the lack of reproducibility is simply due to variations
between experimental repeats or due to the experimental protocol. Thus, the experimental
protocol didn’t seem to work for cefazolin experiments. On the other hand, drug loss results
measured for vancomycin experiments seemed to match with the quantity of drug released by
the scaffolds. Yet again, the loading efficiency of vancomycin calculated for each experimental
repeat was very different. For dexamethasone, depending on the experiment repeat drug loss
results were either matching relatively well the drug release results or were incoherent. Finally,
paclitaxel drug loss results are explained in the next section.
3.3.3 Paclitaxel drug release
Fig. 8 corresponds to one of the three experimental repeats of the paclitaxel drug release
studies. This figure is here to draw attention to the amount of drug released by porous scaffolds
compared to non-porous scaffolds. If different amounts of paclitaxel were released depending
Figure 8. Paclitaxel cumulative release results normalised by the weight of scaffolds. High dose = 10
µg/mg scaffold. Medium dose = 2 µg/mg. Low dose = 0.4 µg/mg. The bars correspond to SD (N=1 / n=6)
P a g e 66 | 78
on the experimental repeat, porous scaffolds released significantly more drug than non-porous
scaffolds in all experimental repeats. When looking at the PCL films on Fig.5, we can observe
that the deposition of paclitaxel seems to be different on the porous and non-porous surface.
Thus, the deposition of paclitaxel on non-porous films is very heterogeneous as we can see
on Fig.5(b), therefore, this first result correlates with the idea that a bigger amount of paclitaxel
could be loaded on porous scaffolds, as observed on Fig.8, thanks to the surface porosity that
seems to allow a better paclitaxel deposition. Moreover, as paclitaxel is deposited in the scaf-
folds pores, the contact surface with the release medium is probably lower than on the non-
porous surface. This hypothesis would imply a faster release of paclitaxel for non-porous scaf-
folds onto which paclitaxel seems to form heterogeneous crystalline patches, yet, as explained
below, no real differences were found between the kinetics of porous and non-porous scaf-
folds. It is important to notice that only a very little amount of paclitaxel has been released
compared to the amount loaded. Fig. 8 we can notice that after 500 hours, porous scaffolds
loaded with high doses of paclitaxel (10µg/mg of scaffolds) released only 1.7µg/mg of the
scaffold, which is only equivalent to 17% of the initial amount of paclitaxel-loaded. Neverthe-
less, the amount of paclitaxel lost in the tube during loading for this experimental repeat was
corresponding to only 26% of the initial amount loaded in average. Suggesting that 57% of the
paclitaxel initially loaded is missing. The same phenomenon happened in the two other exper-
imental repeats. After sonication, a signal equivalent to 5% of the initial amount of paclitaxel-
loaded was measured for both porous and non-porous scaffolds, suggesting that a considera-
ble amount of paclitaxel could be potentially remaining on the surface of both porous and non-
scaffolds after 500 hours. The fact that a considerable amount of paclitaxel is not released
could be explained, either by an interaction between paclitaxel and PCL, or by paclitaxel being
too poorly soluble in PBS/tween 0.1%. If FTIR results aren’t sufficient to disprove entirely the
absence of interaction between PCL and Paclitaxel, it might be more likely that the release
medium is the source of the issue. When normalized by the total amount of drug released (Fig.
9), it is difficult to say if there is a difference between the release kinetics of porous and non-
porous scaffolds. It seems that porous scaffold loaded with a medium dose (Fig 9. b) could
have a very slightly slower release compared to non-porous scaffolds, but this difference is too
small to make any conclusion. Moreover, the results obtained for low dose and high dose are
not suggesting any differences between porous and non-porous scaffolds release (Fig 9. a,
c). The burst release is quite acceptable as after the first 24h only 20% to 30% of the drug is
released. Paclitaxel is then released over the course of approximately 20 days, which is a
significant amount of time, but we can see that after 200 hours the paclitaxel release signifi-
cantly slow down as 80% of the drug content has been released. The fact that release experi-
ments were stopped after 500h for practical reasons also need to be taken into account as a
P a g e 67 | 78
Figure 9. Paclitaxel cumulative release results normalised by the total amount of drug released.
High dose = 10 µg/mg scaffold. Medium dose = 2 µg/mg. Low dose = 0.4 µg/mg. Each point is the
average of 6 replicates (N=3).
(a)
(b)
(c)
P a g e 68 | 78
small amount of paclitaxel was still being released, especially for samples loaded with high
doses. Therefore, technically the release of paclitaxel lasts longer than 500 hours.
3.3.3 Dexamethasone: drug release
When normalized by the total amount of dexamethasone released (Fig 10.), no differences
can be observed in the release kinetic between porous and non-porous scaffolds. Unlike
paclitaxel, the burst release in the first 24h is more important and seems to be dose dependent.
Therefore, after 24h scaffolds loaded with a low dose of dexamethasone released 30% to 65%
of their dexamethasone, while scaffolds loaded with a high dose released 70% to 75%. There-
fore, it seems that the burst release of dexamethasone in the first 24h gets more important as
the dose loaded on the scaffold increase. Finally, most of the dexamethasone is getting re-
leased over the first 4 days. Thus, scaffolds loaded with a low dose of dexamethasone released
80% of their content, while scaffolds loaded with a high dose released 95% of
3.3.4 Dexamethasone: bioactivity
The cell viability of hOB exposed to different doses of free dexamethasone for 24h is presented
in Fig 11. A clear drop in cell viability can be observed between 325 µg/ml and 650 µg/ml of
dexamethasone, statistical analysis revealed that this drop is statistically significant. The mi-
croscopy pictures (data not shown) also shown that in addition to having fewer cells, a signifi-
cant amount have a round morphology when exposed to high dose of dexamethasone. Below
325 µg/ml and above 2.5 µg/ml of dexamethasone, it is difficult to make assumptions about
any effect of dexamethasone, as cell viability remain relatively similar to the cell control. No
differences can be observed either on the light microscopy pictures below 325 µg/ml. The
concentration of 2.5 µg/ml 1 .3 µg/ml seems to slightly increase the cell viability, but no statis-
tically significant differences were found. However, literature tends to suggest that dexame-
thasone is still toxic at a concentration of 1.3 µg/ml[64], consequently, no conclusion can be
made. Fig 12. we can see the cell viability of hOB while exposed to porous and non-porous
scaffolds loaded with three doses of dexamethasone (150µg/scaffold, 30 µg/scaffold, 6
µg/scaffold) over 5 days. After 24h of treatment, it appears that hOB cultured with scaffolds
loaded with high doses of dexamethasone have statistically significantly lower cell viabilities,
while the other conditions are comparable to controls. At day 3, the cell viability of hOB cultured
with dexamethasone-loaded scaffolds have increased compared to day 1 and are statistically
significantly higher than the controls. Interestingly, if no differences were found for porous scaf-
folds; the non-porous scaffolds loaded with a medium and low dose of dexamethasone have
a significantly higher cell viability than the one loaded with a high dose. On the other hand, cell
viability of the controls is lower than at day 1. At day 5, the cell viability of hOB cultured with
P a g e 69 | 78
Figure 10. Dexamethasone cumulative release results normalised by the total amount of drug
released. High dose = 25 µg/mg scaffold. Medium dose = 5 µg/mg. Low dose = 1 µg/mg. Each
point is the average of 6 replicates (N=3).
(a)
(b)
(c)
P a g e 70 | 78
loaded scaffolds is lower than at day 3 but remain higher than at day 1. The control conditions
have similar cell viability than at day 3 and remain statistically significantly lower than scaffolds
loaded with dexamethasone. It is interesting to notice that the standard error means of control
conditions are considerable comparably to other conditions. Compared to day 3, non-porous
scaffolds loaded with medium and low doses-maintained similar cell viabilities, while other
conditions seem to have slightly lower values. Moreover, non-porous scaffolds loaded with
medium and low doses have a statistically significantly higher cell viability than for high doses;
again, no differences were found for porous scaffolds. Consequently, the cytotoxicity associ-
ated with the high dose of dexamethasone suggest the dexamethasone released from the
scaffold is still bioactive. Passed day 1, results suggest that scaffolds loaded with dexame-
thasone increase cell viability similarly to the previous assay. Moreover, for unknown reasons,
Figure 11. Human osteoblasts cell viability assay measured by alamar blue after incubation of
24h with free dexamethasone. Bars correspond to SEM (N=1; n=8). A T-test analysis was car-
ried out to verify the statistical difference of the mean values (**=p < 0.01).
**
**
Figure 12. Human osteoblasts cell viability assay measured by alamar blue at day 1, 3 and 5 when
exposed to scaffolds loaded with dexamethasone (High dose = 150µg/scaffold; Medium =
30µg/scaffold; Low dose = 6µg/scaffold). Bars correspond to SEM (N=3; n=4) A one-way analysis
of variance was carried out to verify the statistical difference of the mean values (**=p < 0.01).
P a g e 71 | 78
Figure 13. Cefazolin cumulative release results normalised by the total amount of drug released.
High dose = 10 µg/mg scaffold. Medium dose = 2 µg/mg. Low dose = 0.4 µg/mg. Each point is the
average of 6 replicates (N=2).
(a)
(b)
(c)
P a g e 72 | 78
this phenomenon seems to be more important for cell exposed to low dose of dexamethasone
on non-porous scaffolds. Yet again literature has shown a decrease of cell viability and an
increase of cell apoptosis of hOB when exposed to a concentration of 10-6M [64]. Thus, it is
difficult to draw a conclusion on the ability of dexamethasone to increase cell viability. One
possible explanation could be that, passed the burst release of day 1, very small amount of
dexamethasone below the concentration of 10-6M are released which then stimulate cell
growth. Finally, it can be noticed that the NaOH leaching process applied to porous and non-
porous scaffolds doesn’t seem to have a negative impact on cell viability, as no statistically
significant differences could be found among the controls.
3.3.5 Cefazolin drug release
Unlike paclitaxel and dexamethasone, when looking at the normalized results Fig 13. we can
clearly see a different release kinetic between porous and non-porous scaffolds. Thus, porous
scaffolds release cefazolin more gradually over more than 100h, while non-porous scaffolds
release all the cefazolin in a burst during the first three hours. Like for dexamethasone, the
burst release of porous scaffolds in the first three hours seems to increase with the dose of
cefazolin. We can also observe that for both porous and non-porous scaffolds, cefazolin seems
to be released faster as the dose increases. Consequently, after 24h, low dose porous and
non-porous scaffolds respectively released 68% and 89% while high dose porous and non-
porous scaffolds respectively released 85% and 99% of the total amount of cefazolin. For un-
known reasons, the experimental protocol to measure drug loss is not working for cefazolin,
and very small values of drug loss are obtained while a significant drug loss is expected as
only 40% of the cefazolin initially loaded was released on average. Finally, those first results
are promising as porous scaffolds seem to be able to release cefazolin over a longer span of
time. However, it is important to take into account the fact that the burst effect and release
kinetics get respectively higher and faster as the dose of cefazolin increases. Consequently, it
could be interesting to see if the difference of release kinetic between porous and non-porous
scaffolds still exist at higher doses.
3.3.6 Vancomycin drug release
We can see Fig 14. results of the vancomycin release experiments normalized by the total
amount of vancomycin released. Even though the difference is not as considerable as the one
observed previously for cefazolin, porous scaffolds are clearly releasing vancomycin slightly
slower than non-porous scaffolds. Interestingly, this difference is hardly visible for low dose
samples. Results of low dose samples are often more spread as the measured values are very
small compared to high doses and are more influenced by measurement variation and errors
P a g e 73 | 78
Figure 14. Vancomycin cumulative release results normalised by the total amount of drug re-
leased. High dose = 10 µg/mg scaffold. Medium dose = 2 µg/mg. Low dose = 0.4 µg/mg. Each
point is the average of 6 replicates (N=3).
(a)
(b)
(c)
P a g e 74 | 78
as such. Unfortunately, the burst release of vancomycin is way more important than for
cefazolin, and most of the vancomycin present on the scaffold surface is released in PBS in
the first three hours for both type of scaffold. It is interesting to notice that similarly to cefazolin
and dexamethasone results, the burst release of vancomycin increases with the dose inde-
pendently of the scaffold type. Therefore, 87% of the vancomycin is released by porous scaf-
fold loaded with high dose in the first hour, while porous scaffold loaded with a low dose re-
leased only 67% of the vancomycin.
3.3.7 Final discussion
To get more insight into the factors influencing the drug release studies, it is important to dis-
cuss the results all together besides analyzing them separately. One main question arises from
this work; why the release kinetics is different between porous and non-porous scaffolds only
for cefazolin and vancomycin? In an attempt to explain the results, it is logical to look first at
the properties of the drug. Cefazolin and vancomycin have in common a high solubility in water,
on contrary to paclitaxel and dexamethasone which both have a very poor solubility in water.
This difference of solubility in water is the first element of the answer, even if it doesn’t explain
by itself how it influences the release kinetics of the drug between a porous and non-porous
scaffold. As mentioned earlier for Paclitaxel, it would be interesting to investigate if a different
release kinetics is observed with a release media in which paclitaxel and dexamethasone are
more soluble. Secondly, it appeared during experimentation that both porous and non-porous
scaffolds are sinking when placed in ethanol during drug loading. Yet, non-porous scaffolds
sink while porous scaffolds float when immersed in PBS or in cell media. This phenomenon
could be explained by the lower surface tension of ethanol combined with the hydrophobicity
of PCL. Thus, small air bubbles could be entrapped in the micropores of porous scaffolds,
making the scaffolds float in PBS but sink in liquid with lower surface tension such as ethanol.
The presence of air bubbles on the scaffold surface could also reduce the surface contact
between the release medium and the scaffold, which as a result would reduce the amount of
drug released. Hence, drug release results might be different if porous scaffolds were fully
immersed in the release media. Further experimentations could also help us have a better
understanding of what’s happening. Therefore, it would be important to extend the FTIR anal-
ysis to cefazolin, vancomycin, and dexamethasone as the difference between the results could
also due to a possible interaction between the drug and the scaffold. Indeed, depending on the
type of bonding created between the drug and the scaffold after loading, the release of the
drug could be hindered if immersion in the release media is not sufficient to detach the drug
from the scaffold. Moreover, it could also be interesting to analyze the homogeneity of each
drug deposition on the surface of porous and non-porous scaffolds as the release will directly
depend on the way each drug adheres to the surface. It was clear from Fig.5 that deposition
P a g e 75 | 78
of paclitaxel on non-porous films was not heterogeneous while no conclusion could be made
for the porous film as the limitations of light microscopy, notably the light reflection of PCL,
don’t allow us to visualize drug deposition in the scaffold pores. However, more advanced
imaging technologies such as scanning electronic microscopy might make possible the obser-
vation of the drug deposition in the pores.
4. CONCLUSION AND FUTURE WORK
First, this work investigated the printability of a novel PCL/PBS composite. Rheological meas-
urements have shown that as the mass ratio of PBS microparticles increase PCL is losing its
linear viscoelastic region, in addition, to being less sensitive to temperature increase. The
PCL/PBS composite was then successfully printed into scaffolds with a screw-based extrusion
3D printing technique by increasing the manufacturing temperature. Secondly, this work inves-
tigated the potential of PCL scaffolds as a drug carrier, and more precisely how the scaffold
microporosity is influencing the release kinetics of different drugs. Our results have shown that
porosity seems to decrease the burst effect and prolong the drug release of soluble drugs
compared to non-porous scaffolds. On the other hand, no significant differences could be ob-
served for scaffolds loaded with insoluble drugs. But despite some encouraging results, the
burst release remains important for soluble drugs, and as such, their clinical relevance remains
limited. Concerning the influence of the loading process, results seem to indicate that it is not
affecting the drug bioactivity since a high concentration of dexamethasone has shown a nox-
ious effect on cell viability. Obviously, further work is required to verify the drug bioactivity of
Paclitaxel, cefazolin, and vancomycin. Finally, results of drug release studies have also shown
a lack of control over the drug loading process. Indeed, the total amount of drug released
between experimental repeats was considerably different. As such, there is a strong need to
find an assay that would allow us to measure accurately the drug loss during the loading pro-
cess. Without this data, it is difficult to draw a conclusion as we’re unable to know precisely
how much drug has been released comparatively to the total amount loaded on each scaffold,
which would maybe show different release kinetics. To follow up this work two experiments
could be carried out to shed more light on the results. Firstly, HPLC could maybe be used to
measure the amount of drug left in the tube after loading; secondly, the scaffolds could maybe
be degraded in a solution which doesn’t affect the drug to measure how much drug is left on
the scaffolds at the end of the drug release study. Thus, more work is required to determine if
experimental conditions are optimal to get more insight into how porosity exactly affects drug
loading and drug release.
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