encapsulation of valvular microtissues in modified gelatin...encapsulation of valvular microtissues...
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Encapsulation of valvular microtissues
in modified gelatin
Word count: 19,172
Yohana Permatasari Handoyo Student number: 01101853
Supervisor(s): Prof. dr. Ria Cornelissen
Mentor : Annelies Roosens
A dissertation submitted to Ghent University in partial fulfilment of the requirements for the degree of
Master of Science in the Biomedical Sciences
Academic year: 2016 – 2017
Encapsulation of valvular microtissues
in modified gelatin
Word count: 19,172
Yohana Permatasari Handoyo Student number: 01101853
Supervisor(s): Prof. dr. Ria Cornelissen
Mentor : Annelies Roosens
A dissertation submitted to Ghent University in partial fulfilment of the requirements for the degree of
Master of Science in the Biomedical Sciences
Academic year: 2016 – 2017
ACKNOWLEDGEMENT
There are so many people to thank for their help, support, guidance and efforts during this
master dissertation. I could never have accomplished the heights or explored the depths
without them.
My sincere gratitude goes first to Prof. dr. Ria Cornelissen, my supervisor, who has given me
the opportunity to join Tissue Engineering Research Department where I had challenging yet
valuable experiences to complete this master dissertation and where I have picked up a lot of
knowledge and skills during the past two years.
A very special thank you to my mentor, Annelies Roosens, for her incredible patience and
enthusiasm. I have learned a lot throughout these years and I am inspired by her
professionalism and by the way she executes plans and solves problems. Thank you for
always being open to my questions and ideas, and for always challenging and motivating me
in developing high quality work.
I also would like to express my gratitude to the whole department and all staff members,
especially Leen, Greet, and Johanna, for the practical experiences and training, and also to
my fellow students, for the entertaining chitchat in our free time.
Personally, I am deeply grateful to my uncle, who has offered me the chance to realize one of
my biggest dreams, which is studying abroad. Without him my dream would have just been a
sweet dream. Thank you for the encouragement and all kinds of support during my study and
my stay in Belgium.
To my sisters, Jessica and Josephine, I would like to thank them for constantly listening to my
stories and for cheering me through the difficult moments. Mama and Papa, although you are
not physically present here, your never-ending support and prayers are always with me.
Last but not least, my heartfelt thanks to my fiancé, Klemens, who remains willing to provide
me with his unfailing encouragement from miles away and for understanding the less frequent
contact in these two busy years.
Above all, I believe that I could have done all these things through Christ who strengthens me,
and by His Grace I have been surrounded by these kind-hearted people.
Yohana Permatasari Handoyo
TABLE OF CONTENTS
SAMENVATTING .................................................................................................................. 1
SUMMARY ............................................................................................................................ 2
INTRODUCTION ................................................................................................................... 3
1. Background ................................................................................................................ 3
2. Aortic valve ................................................................................................................. 3
2.1. Anatomy (macrostructure) ................................................................................... 3
2.2. Histology (microstructure) .................................................................................... 4
3. Aortic valve pathologies .............................................................................................. 6
4. Heart valve replacements ........................................................................................... 7
4.1. Mechanical heart valves ...................................................................................... 7
4.2. Bioprosthetic heart valves .................................................................................... 8
5. Heart valve tissue engineering (HVTE) ....................................................................... 9
5.1. Top-down tissue engineering (TD-TE) ................................................................. 9
5.2. Bottom-up tissue engineering (BU-TE) ................................................................ 9
6. Hydrogels ..................................................................................................................11
6.1. Crosslinking of hydrogels ....................................................................................11
6.2. Hydrogels for HVTE ............................................................................................12
7. 3D (bio)printing ..........................................................................................................12
8. Goal of the master thesis ...........................................................................................13
MATERIALS AND METHODS ..............................................................................................14
1. Cell isolation ..............................................................................................................14
2. 2D culture ..................................................................................................................14
3. 3D culture ..................................................................................................................14
3.1. Agarose microwell chips .....................................................................................14
3.2. Micro-aggregate formation ..................................................................................15
3.3. Micro-aggregate collection ..................................................................................15
3.4. Methacrylamide-modified gelatin (Gel-MOD) synthesis .......................................15
3.5. Cell/micro-aggregate encapsulation in Gel-MOD ................................................16
3.6. Fusion of micro-aggregates in Gel-MOD .............................................................17
4. Viability assay ............................................................................................................17
5. Histochemical staining ...............................................................................................17
5.1. Hematoxylin-eosin staining .................................................................................17
5.2. Immunohistochemical staining ............................................................................18
6. RT-qPCR ...................................................................................................................19
7. Gel-MOD characterization .........................................................................................20
7.1.1. Equilibrium swelling test ..................................................................................20
7.1.2. Compression test ............................................................................................21
8. Data analysis .............................................................................................................21
RESULTS .............................................................................................................................22
1. Encapsulated single cells in Gel-MOD .......................................................................22
2. VIC micro-aggregates ................................................................................................23
3. Encapsulated VIC micro-aggregates ..........................................................................25
4. Gene expression of encapsulated VIC micro-aggregates ..........................................31
5. Fusion of VIC micro-aggregates in Gel-MOD .............................................................34
6. Gel-MOD characterization .........................................................................................35
DISCUSSION .......................................................................................................................38
CONCLUSION .....................................................................................................................44
FUTURE PERSPECTIVES...................................................................................................45
REFERENCES .....................................................................................................................46
SUPPLEMENT 1 ..................................................................................................................... i
SUPPLEMENT 2 ................................................................................................................... iii
SUPPLEMENT 3 ................................................................................................................. viii
1
SAMENVATTING
Deze studie onderzocht of methacrylamide-gemodificeerde gelatine (Gel-MOD) een geschikt
inkapsel-materiaal is voor single cells/micro-aggregaten van valulaire interstitiële cellen (VIC),
om een bio-inkt te verkrijgen voor de biofabricatie van een groter valvulair macroweefsel. VIC
single cells en micro-aggregaten werden ingekapseld in 10 en 20 w/v% Gel-MOD en
gecrosslinkte in de aanwezigheid van 2 mol% Irgacure 2959 en 20 mol% VA-086.
Deze studie heeft aangetoond dat de viabiliteit van VIC single cells snel afnam na inkapseling,
terwijl de ingekapselde micro-aggregaten viabel bleven gedurende de volledige opvolgperiode.
De ingekapselde micro-aggregaten waren ook in staat om hartklep-gerelateerde ECM
componenten te produceren en te remodelleren.
De Gel-MOD concentratie en type foto-initiator beïnvloedde de gelstijfheid, maar hadden geen
invloed op de viabiliteit van aggregaten. In de zachtste 10 w/v% VA-gecrosslinkte gels
vertoonden VIC micro-aggregaten meer uitgesproken proliferatie, migratie en waren in staat
om te fuseren en grotere valvulaire microweefsels te vormen, maar deze gels waren minder
stabiel dan de stijvere Irgacure-gecrosslinkte gels.
VIC micro-aggregaten in de stijvere gels vertoonden downregulatie van α-SMA- en upregulatie
van FSP-1 expressie, wat erop wees dat VIC zich in een homeostatische quiescent toestand
bevonden. De GAG-productie in deze gels was opmerkelijk gestegen en de vroege
chondrogene merker, Sox-9, was significant verhoogd.
Onze resultaten toonden aan dat Gel-MOD kan worden gebruikt voor het inkapselen van VIC
aggregaten en dat zachter gels meer geschikt zijn dan stijvere gels voor de generatie van een
groter valvulair macroweefsel. De stabiliteit van deze zachtere gels moet echter nog verbeterd
worden en er dient ook nog bepaald te worden of VIC in een homeostatische quiescent,
ongedifferentieerde toestand kunnen blijven in soft-gel ingekapselde microaggregaten.
2
SUMMARY
This study investigated whether methacrylamide-modified gelatin (Gel-MOD) is a suitable
carrier-material for valvular interstitial cells (VIC) single cells/micro-aggregates, to obtain a bio-
ink for the biofabrication of a larger macro-scale valvular tissue. VIC single cells and micro-
aggregates were encapsulated in 10 and 20 w/v% Gel-MOD and crosslinking was performed
in the presence 2 mol% Irgacure 2959 and 20 mol% VA-086.
This study has shown that the viability of VIC single cells decreased rapidly after encapsulation,
while the encapsulated micro-aggregates remained viable during the whole study duration.
The encapsulated micro-aggregates were also able to produce and remodel valve-related
ECM components.
Gel-MOD concentration and photo-initiator type influenced the gel stiffness but had no impact
on aggregate viability. In the softest 10 w/v% VA-crosslinked gels, VIC micro-aggregates
showed more pronounced proliferation, migration, and were capable to fuse and to form larger
valvular microtissues, but these gels were less stable than the stiffer Irgacure-crosslinked gels.
VIC micro-aggregates in the stiffer gels showed downregulation of α-SMA- and upregulation
of FSP-1 expression, indicating that VIC remained in a homeostatic quiescent state. However,
GAG production in these gels was extremely increased and the early chondrogenic marker,
Sox-9, was significantly elevated.
Our data showed that Gel-MOD can be used for the encapsulation of VIC aggregates and
indicated that softer gels are more suitable than stiffer gels for the generation of a macro-scale
valvular tissue. However, the stability of these softer gels needs to be improved and it needs
to be determined if VIC can remain in a homeostatic quiescent, undifferentiated state in soft-
gel encapsulated micro-aggregates.
3
INTRODUCTION
1. Background
Heart valves play a key role in the unidirectional blood flow regulation of the body. The mitral
and tricuspid valve are atrioventricular valves located between the atria and the ventricles,
while the pulmonary and aortic valve are semilunar valves that allow the ventricular outflow to
the pulmonary artery and to the aorta respectively [1].
Heart valve diseases are believed to be ‘the next cardiac epidemic’ [2]. An estimated 2.5% of
the population in developed countries suffer from heart valve diseases. One of the most
commonly affected and transplanted valves is the aortic valve [3].
2. Aortic valve
2.1. Anatomy (macrostructure)
The aortic valve is located in the transition part from the left ventricle to the ascending aorta.
This transition can be seen at the change of the muscular layer on the ventricular side to the
fibro-elastic layer of the atrial wall. The aortic valve consists of three semilunar leaflets and the
aortic root which includes the annulus or ventriculo-aortic junction, the leaflet attachment site,
the three sinuses of Valsava, three interleaflet triangles, three commissures, and the
sinotubular junction (Fig. 1) [4-6].
Figure 1: Schematic overview of the aortic valve and root components [5]
The annulus or ventriculo-aortic junction is the first transition part and the attachment site of
the valvular leaflets. The leaflet attachment site is U-shaped and has a crown-like structure
that plays a role as the hinges of the semilunar leaflets [4-6]. The sinuses of Valsava, of which
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two sinuses are responsible for the coronary arteries perfusion, consist of three bulges at the
aortic walls. These structures are situated between the annulus and the sinotubular junction
[5, 6]. The sinotubular junction is a circular ring at the top of Sinuses of Valsava and is the
boundary between the aortic root and the ascending aorta. [4-6]. The commissures are located
at the apical part of the leaflet attachment site, above the interleaflet triangles [5, 6]. The aortic
valve has three semilunar leaflets, of which each has three distinct parts; the free margin or
coaptation area of the leaflet to the neighboring leaflets, the belly part, and the basal
attachment part [5]. Aortic valve leaflets are avascular tissues that receive nutrients through
diffusion and hemodynamic convection [7].
2.2. Histology (microstructure)
The extracellular matrix (ECM) components of the aortic valve leaflets, such as type I, III and
V collagens (Col I, III, and V), glycosaminoglycans (GAG), and elastin, give a major
contribution to the biomechanical characteristics of the leaflets [7-9]. The ECM is divided into
three distinct layers; the lamina fibrosa (LF) at the atrial side, the lamina spongiosa (LS), and
the lamina ventricularis (LV) at the ventricular side (Fig. 2A). The overall proportion of collagen,
proteoglycans (PG), and elastin in the human heart valves is respectively 60%, 30%, and 10%
[9]. Collagens are situated in all layers of the leaflets. However, circumferentially aligned Col I
is one of the main ECM components predominantly found in the LF [7, 9]. The collagen bundles
are responsible for the stiffness of the leaflets by applying resistance against the diastolic
tension during the closed valve condition (Fig. 2B) [7]. PG consist of one or more linear
negatively charged polysaccharide GAG chain(s), covalently attached to a serine residue of a
core protein [8]. PG are abundantly present in the middle layer (LS) [7, 9] and are hydrophilic
molecules that are capable to compensate the shearing as a result of LF and LV deformation
by absorbing water [8, 10]. The major PG in valve tissue are biglycan, decorin, and versican.
These PG contain chondroitin sulfate and/or dermatan sulfate GAG side chains [11]. Total
GAG composition in the valve is approximately 30% chondroitin-4-sulfate/chondroitin-6-
sulfate, 15% dermatan sulfate, and 55% hyaluronan (HA) [12]. HA is the only GAG that does
not bind to a core protein and is exclusively non-sulfated [13]. Radially-oriented elastin
comprises mainly the LV and provides elasticity of the valve leaflets especially during the
systole by stretching the elastic fibers (Fig. 2B) [7, 8].
The surfaces of the leaflets are protected by a monolayer of valvular endothelial cells (VEC)
(Fig. 2A). This monolayer plays a significant role in thrombogenesis prevention during blood
contact, nutrient supply, as well as biochemical transmission to the underlying VIC. VEC
behave differently based on their location on the leaflets. VEC on the aortic side have more
5
tendency to protect the leaflets against lesion and inflammation since the shear stress at the
fibrosa side is higher than the ventricularis side [7].
The ECM is maintained and synthesized by valvular interstitial cells (VIC) that are distributed
throughout all the matrix layers and play a significant role in the functioning of heart valve
leaflets [14]. During embryonic development, VIC arise from the transformation of the
endothelial cells on the endocardial cushion, which is known as the endothelial-mesenchymal
transition (EMT) [1, 14]. VIC can also arise from progenitor VIC (pVIC) from bone marrow (Fig.
3). pVIC have a high proliferative capacity which is important to repair an injured heart valve.
However, the availability of pVIC is very limited and further studies on the contribution of these
cells during valve repair are necessary [14].
Figure 2: A. The three layers of a heart valve leaflet B. Role of elastin and collagen during leaflet deformation (A) The aortic valve leaflets consist of three ECM layers: the lamina fibrosa, spongiosa, and ventricularis, and is lined with valvular endothelial cells (adapted from [15]). (B) During the systole stage collagen recoils and elastin crimps, while during the diastole collagen tenses and elastin stretches (adapted from [7]).
VIC are a heterogeneous population and can be divided in subpopulations (Fig. 3) [7]. In the
normal aortic valve, VIC are in a quiescent, low proliferative state (qVIC) [1, 14, 16]. qVIC keep
the aortic valve avascular by the secreting anti-angiogenic factor, chondromodulin-1 [17]. In
addition, qVIC have fibroblastic features and express high vimentin (vim) and low α-smooth
muscle actin (α-SMA) [16]. Under certain pathophysiological conditions, VIC become activated
(aVIC) and responsible for the remodeling of the aortic valve. This activation is characterized
by high cell proliferation and migration, along with an increase of matrix metalloproteinases
(MMP) and tissue inhibitor of metalloproteinases (TIMP) that respectively induce ECM
secretion and ECM degradation. aVIC are myofybroblast-like cells and marked by high
expression of α-SMA and low expression of vim [1, 14, 16]. The aVIC can lead to heart valve
calcification via a dystrophic/apoptotic mechanism. On the other hand, osteoblast-like VIC
(obVIC) play a major role in the calcification of the aortic valve via an osteogenic pathway.
Some chondrogenesis and osteogenesis-associated genes, such as Runx2, Sox9 [18], early
A B
6
growth factor (Egr-1) [19], osteocalcin (OCN), and osteopontin [20] are upregulated in calcified
heart valve. To date, there is still no strong evidence whether obVIC arise directly from qVIC
or through a precedential stage of aVIC [1, 14, 21].
Figure 3: Scheme of VIC phenotype conversion (PhD thesis Roosens A.)
3. Aortic valve pathologies
Aortic valve diseases (0.9%) bring a certain contribution to the total heart valve diseases (the
age-adjusted prevalence of valve diseases is 2.5% [3]). The two most prevalent valve diseases
are aortic stenosis (AS) (0.4%) and aortic insufficiency (AI) (0.5%) [22].
AS refers to the thickening of aortic valves which results in the narrowing of aortic valve
opening and in abnormal leaflet motion. This condition leads to left ventricular hypertrophy due
to progressive high pressure overload [23, 24]. AS can also be caused by severe calcification
and valve degeneration, which mostly occur in elderlies. However, AS may also be present in
young children due to congenital bicuspid aortic valves, rheumatic fever, and endocarditis [2,
7, 24].
AI, also known as aortic regurgitation, is mainly caused by bicuspid aortic heart valves and
aortic root dilatation [2]. Nevertheless, AI may also appear as a result from other factors similar
to AS. Due to this pathology, the aortic valve closes incompletely and leaks during the diastole.
This results in the backward blood flow from the aorta to the left ventricle whereby the heart
needs to perform more effort to supply sufficient blood [25].
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4. Heart valve replacements
Dysfunctional heart valves are first attempted to be repaired by various surgical techniques.
However, when there is no possibility to repair the valve, the standard approach is to replace
the aortic valve with a mechanical or bio-prosthetic valve [26, 27]. Heart valve prostheses
should possess excellent durability, a non-thrombogenic surface, appropriate hemodynamic
performance, and implantability [27].
Figure 4: Heart valve prostheses [27]. 1. Mechanical valves (A-C). (A) Bileaflet valve (B) Tilting-disc valve (C) Caged-ball valves. 2. Bioprosthetic valves (D-H). (D) Stented porcine valve (E) Stented bovine valve (F) Stentless porcine valve (G) and (H) Transcatheter bioprosthesis
4.1. Mechanical heart valves
Mechanical heart valves (MHV) are manufactured of metallic and ceramic biomaterials. In
general, MHV consist of hinges, a stent, leaflets and a sewing ring, and they are developed in
several designs, such as caged-ball valves, tilting-disc valves, and bileaflet valves (Fig. 4A-C)
[27, 28].
The first implantable caged-ball valves, Starr-Edwards were made in 1960. The durability of
these prostheses have been proven. Nevertheless, caged-ball valves cause some health
problems due to high pressure drop and inappropriate hemodynamics (Fig. 4C) [28]. The next
developed MHV were tilting-discs which consist of a single pyrolitic-carbon disc fastened by
metal struts (Fig. 4B) [27, 28]. Although these MHV have less hemodynamic problems than
caged-ball valves, stent fractures and other complications are still present [28]. Bileaflet valves
are the leading design of MHV with a similar concept to the tilting disc valves. These
prostheses have a pair of semicircular discs that are anchored to the valve ring by hinges (Fig.
4A) [27, 28]. Compared to the precedent MHV, these prostheses do not disturb the normal
laminar blood flow. The newer-generation valves require less anti-coagulation in comparison
with the older models. Nevertheless, thromboembolisms may still occur, especially around the
hinge area of the valve due to shear stress of the blood flow. Therefore, patients with MHV still
8
require a life-long anti-coagulation therapy [28]. The administration of anti-coagulant, mostly
warfarin, is not without risks (e.g. hemorrhagic risk) [29]. Moreover, warfarin has teratogenic
effects, which appears to be a contraindication for young women who still expect pregnancies
[26].
MHV are durable and can function well for 20-30 years [29]. MHV are primarily addressed to
patients with less bleeding issues, especially to older patients (but still under 65 years old) with
a long life expectancy with purpose to avoid open heart re-surgery [27, 28].
4.2. Bioprosthetic heart valves
Bioprosthetic heart valves (BHV) (Fig. 4D-H) are an alternative to MHV as BHV do not require
the patients to receive anti-coagulation therapy. Therefore, BHV are addressed to patients
above 65 years old who have shorter life expectancy and higher bleeding risk [27, 28].
Furthermore, these prostheses are also implanted in younger or pediatric patients to avoid the
side effects of a life-long anti-coagulation therapy [28]. BHV have a limited durability due to
valve degeneration and calcification, and re-operation will be necessary after 10-15 years [29].
BHV can be obtained from auto-/allograft (human origin) and xenograft (animal-origin) [27, 28].
4.2.1. Xenografts (heterografts)
BHV can be obtained from porcine aortic valves or bovine pericardia [27]. These materials are
chemically crosslinked with glutaraldehyde (GA). This pretreatment aims to fix the tissues in
order to avoid autolysis and to improve mechanical stability. GA also kills the resident cells
which reduces the immune reaction of the host to the implanted BHV [28, 30, 31]. During
manufacturing, xenograft valves can be attached on a metal loop (stented) (Fig. 4D&E) or
without a supporting material (stentless) (Fig. 4F). Recently, BHV can also be implanted
percutaneously via transcatheter aortic-valve implantation (TAVI) (Fig. 4G&H) to avoid the
high surgical risk. This minimal invasive technique is suitable for patients who are inoperable
due to their age, left ventricular dysfunction, and other health conditions [32].
4.2.2. Auto-/allograft (homograft)
Diseased heart valves can also be replaced by donor heart valves (allografts). Unlike
xenografts, the donor heart valves are cryopreserved in liquid nitrogen before implantation.
Allografts have a minimal infection risk and apply good mechanical functions. However, the
availability of allograft valves is limited [30, 31].
Autografts refer to heart valves that are taken from the patients themselves. The replacement
of diseased aortic valves in pediatric patients is commonly done based on the Ross procedure.
In this surgery, the aortic valve is replaced by the patient’s own pulmonary valve and in place
of the pulmonary valve, an allograft is implanted [26, 30]. The most significant advantage of
this procedure is that patients receive a viable aortic valve replacement which has capacity to
9
grow together with the growth of the child. However, there is high mortality risk in young-aged
patients during the surgery [26].
5. Heart valve tissue engineering (HVTE)
Heart valve prostheses are artificial devices that have no ability to grow or to remodel. To
overcome these limitations, HVTE is proposed as a novel method to generate viable aortic
valves which resemble the structure and the functions of the native valve. In TE there are
currently two main approaches, the top-down and the bottom-up approach [15, 33].
5.1. Top-down tissue engineering (TD-TE)
In the top-down approach, cells are seeded onto the surface of a porous scaffold and expected
to penetrate and distribute in the scaffold (Fig. 5A). The scaffolds used in this approach are
bioresorbable/biodegradable scaffolds or acellular tissue scaffolds [34-36].
Bioresorbable/biodegradable scaffolds can be made of synthetic or natural-derived polymers.
The advantages of synthetic polymers are the minimization of infection risks, the reduction of
immunogenic issues, and the possibility to tune the mechanical features and the degradation
time of the scaffold. However, the synthetic polymers do not easily interact with the host body
due to the lack of cell recognition signals. For this reason, modifications of this scaffold, such
as adhesion peptides incorporation, are necessary to improve the physical and chemical
properties [37, 38]. The most commonly used synthetic polymers in HVTE are poly(glycolic
acid) (PGA), poly(lactic acid) (PLA), and poly(lactic-co-glycolic acid) (PLGA). The synthetic
scaffold can be generated by electrospinning, phase separation, freeze drying, and self-
assembly [34, 37]. On the other hand, the natural-derived polymers can mimic the natural
mechanical and physiological properties. The natural-derived polymers comprise polypeptide,
polysaccharide, nucleic acid, and ECM components of tissue such as collagen, fibrin, and HA.
Most of these ECM components are used to make hydrogel scaffolds and will be further
discussed in section 6.2. Hydrogels for HVTE [38-40].
Acellular tissue scaffolds can be generated from allogeneic or xenogeneic tissues (heart
valves, pericardium, and small intestinal mucosa) that are treated by chemical or physical
decellularization techniques to remove the donor cells in order to eliminate the antigenic
epitopes and intracellular components [41, 42].
5.2. Bottom-up tissue engineering (BU-TE)
Due to the restriction of TD-TE to control cell density, diffusion of nutrients, and sometimes
vascularization of the tissue construct, a bottom-up approach is proposed in tissue
engineering. In BU-TE, a modular tissue is built of micro-scale building blocks (cell aggregates,
cell sheets, and cell/aggregates-laden hydrogels) via different assembly techniques, such as
sticking, magnetic-, random-, and directed assembly (Fig. 5B) [43].
10
3D cultures play an important role in BU-TE since 3D cultures more closely reflect the in vivo
micro-environment than the conventional 2D cultures. In a 3D culture, cells have the possibility
to anchor to the neighboring cells and ECM components as they do in the native condition.
These interactions influence cell morphology and facilitate cellular signaling, which has an
effect on cell behavior including cell survival, cell differentiation, cell functionality, and cell
migration [44, 45].
Figure 5: A. Top-down tissue engineering B. Bottom-up tissue engineering. (A) In TD-TE, cells are seeded onto a scaffold and are expected to eventually repopulate the scaffold and produce ECM. (B) In BU-TE, microscopic microtissues are assembled to form tissue constructs via various techniques (Adapted from [43])
There are two different concepts of 3D culture methods: scaffold-based and scaffold-free
cultures. In scaffold-based cultures, the scaffold plays a role as the ECM and as a platform for
cell growth [44]. For instance, cells are cultivated via encapsulation in gels [46]. In scaffold-
free procedures, cells will assemble themselves in 3D micro-aggregates/spheroids [44].
There are many techniques to generate 3D aggregates without using a scaffold, including
pellet culture, spinner culture, hanging drop, liquid overlay, rotating well vessel, external force,
cell sheets, and microfluidics. However, all these methods have limitations such as size control,
shear stress, number of yield, etc. [47, 48]. Recently, a high-throughput technique is
developed, which generates micro-aggregates by using agarose micro-wells. This technique
11
allows to produce aggregates in high quantities and with uniform shape and dimensions in a
low shear force environment [49].
6. Hydrogels
Hydrogels are biocompatible materials with a hydrophilic character and a three dimensional
macromolecular network that are popularly used in tissue engineering. Due to the
hydrophilicity, hydrogels are able to absorb water. Moreover, hydrogels can be degraded,
disintegrated, and dissolved [50-52]. Hydrogels can be derived from both synthetic and natural
polymers. Synthetic hydrogels ease the users to control the mechanical, physical and chemical
properties. In some cases, these hydrogels are modified in order to enhance the biological
function of the synthetic hydrogels and to facilitate the interaction with cells. In contrast to
synthetic hydrogels, natural hydrogels, including collagen, HA, alginate and chitosan, can
support cellular functions and have similar features to the natural ECM. However, modification
of natural hydrogels is necessary to increase the mechanical properties [51, 53]. In TE,
hydrogels are used not only to produce porous scaffolds (TD-TE), but also to encapsulate cells
(3D culture, bio-printing: BU-TE) [52].
6.1. Crosslinking of hydrogels
Crosslinking is essential to create a stable 3D cell-laden hydrogel construct [50]. The type of
crosslinking is dependent on the character of the hydrogels [54]. Hydrogels can be either
physically or chemically crosslinked. Physical crosslinking is a reversible process that is
obtained due to molecular entanglement and secondary forces such as ionic, H-bonding, or
hydrophobic bonding, while chemical crosslinking is an irreversible process due to covalent
bonding. The latter permits the control of the crosslinking density and the degradation rate [50,
54].
By covalently crosslinking, the swelling characteristic of the hydrogels is limited, but the
materials can still be degraded. Chemical crosslinking can be obtained by using chemical
crosslinkers, such as glutaraldehyde, formaldehyde, and carbodiimide [53]. The concern that
remains in the use of crosslinkers is the interaction of these solvents with the proteins in the
native environment that can cause negative effects [55]. Another method is the incorporation
of photo-crosslinkable polymer chains which can be chemically crosslinked in the presence of
a radical photo-initiator (PI) that will react with the target molecules after UV irradiation.
Nevertheless, the toxicity of the radical agent must also be taken into account in the selection
of the PI [40, 53]. The most commonly used PI in tissue engineering is Irgacure 2959
(𝜆=275nm) [56]. However, other PI, such as VA-086 (𝜆=375nm), which have less cytotoxic
effects, are sometimes used [57, 58].
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6.2. Hydrogels for HVTE
The most commonly used hydrogels in HVTE are methacrylamide-modified gelatin (Gel-MOD),
methacrylamide-modified hyaluronic acid (HA-MA), and poly(ethylene glycol) diacrylate (PEG-
DA) [59, 60].
PEG is a synthetic hydrogel which has a hydrophilic character and is largely used in TE due
to the biocompatibility and non-immunogenicity. Moreover, the mechanical properties and the
degradability of PEG can be controlled. To induce biological characteristics in PEG, PEG can
be modified with bioactive molecules, such as adhesive peptides and enzyme-degradable
peptide sequences [40, 52].
HA is one of the ECM components that is present in most tissues and accounts for 50% of the
total GAG in the heart valve leaflets. HA can be naturally degraded by hyaluronidase and
regulates the cell mobility and adhesion [40, 52].
Gelatin is a denatured product of native triple helix collagen. This hydrogel is produced by
acidic (type A gelatin) or alkaline treatment (type B gelatin) of Col I [51]. Gelatin has been
recognized as a safe material by the U.S Food and Drug Administration and has similar
structural and mechanical characteristics to natural collagen, which is the main ECM
component of the heart valve leaflets [52, 60].
7. 3D (bio)printing
In TE, 3D (bio)printing has gained great interest since heart valves can be rapidly
manufactured with this technique. This technique allows to produce patient-specific organs by
transforming patient’s images, which are captured by medical imaging equipment, into a 3D
construct. 3D printing can be used in TD-TE to manufacture scaffolds [61], bioprinting on the
other hand, can be used in BU-TE. In this BU approach, 3D bioprinting is used for the directed
assembly of micro-scale building blocks by layer-by-layer positioning of bio-ink, which may
consist of cell/aggregates-laden hydrogels [62]. This approach offers many advantages related
to the accuracy of the heart valve architecture, the possibility to create mechanically
heterogeneous heart valve, and the control of the cell distribution [63, 64].
3D bioprinters are categorized into three groups; inkjet-based, laser/light based, and extrusion-
based printing (Fig. 6). Inkjet-based printing is the most commonly used technique. The
cartridge is filled with bio-ink [63] and droplets of this bio-ink are forced out of the nozzle based
on a temperature system (thermal inkjet), a piezoelectric actuator, or an ultrasound pressure
system (acoustic inkjet) [64]. In extrusion-based bioprinting, air pressure (pneumatic and
piston) and
motor forces (screw) are utilized to continuously extrude bio-ink from the nozzle. The third type
is laser/light-based bioprinting, which applies laser pulses on an energy-absorbing layer that
13
is able to transfer the applied energy to the donor slide coated with the bio-ink. This energy will
induce the droplets of the bio-ink. [63-65].
Figure 6: Three categories of bioprinting. (A) Inkjet bioprinting (B) Extrusion bioprinting (C) Laser-assisted bioprinting [64]
8. Goal of the master thesis
The aim of this master thesis is to evaluate Gel-MOD hydrogels as VIC aggregate-carrier
material with the purpose of using this bio-ink for the 3D bioprinting of larger scale tissue
engineered valvular tissue in future experiments.
In this study, VIC micro-aggregates will be generated and encapsulated in different Gel-MOD
concentrations. Small hydrogel discs will be formed and hydrogels will be crosslinked with UV
light, in the presence of Irgacure 2959 or VA-086. To investigate if Gel-MOD is a suitable
aggregate-carrier material, the viability, proliferation, ECM production and phenotype of the
encapsulated VIC aggregates will be analyzed (Fig. 7).
Figure 7: Schematic overview of the experimental outline of this master thesis
14
MATERIALS AND METHODS
1. Cell isolation
VIC were isolated from fresh porcine hearts that were obtained from a local slaughter house
in (Ruyckaert N.V., Eeklo). Aortic HVL were dissected and rinsed in Phosphate Buffer Saline
(PBS), supplemented with antibiotica (Penicillin/Streptomycin; Sigma-Aldrich®, Bornem,
Belgium). The leaflets were incubated in an enzyme-mix (Table 1.1) at 37°C for 30min.
Afterwards, the leaflets were manually scraped to remove the VEC and the cell suspension
was centrifuged at 1000rpm for 5 min. The HVL were then treated with another enzyme-mix
(Table 1.2) for approximately 4 hours at 37°C. The cell suspension was filtered by a 70µm cell-
strainer and centrifuged at 1000rpm for 5 min. Afterwards, the VIC were seeded in T175
polystyrene flasks filled with complete DMEM (cDMEM+) (Table 1.3).
Table 1: enzyme-mix and culture media for VIC isolation and cultivation 1. Enzyme-mix for VEC removal
Total Volume 50 ml
DMEM Glutamax (Gibco®, Life TechnologiesTM) 50 ml
0,1% type 1 collagenase (17100-017, InvitrogenTM, Merelbeke) 0.05g
0,12% dispase (D4693, Sigma) 0.06g
2. Enzyme mix for VIC isolation
Total Volume 50 ml
DMEM Glutamax (Gibco®, Life TechnologiesTM) 50 ml
0,1% type 1 collagenase (17100-017, InvitrogenTM, Merelbeke) 0.05g
3. cDMEM+ for VIC
Total Volume 500 ml
DMEM Glutamax (Gibco®, Life TechnologiesTM) 500ml
10% Fetal Calf Serum (InvitrogenTM, Merelbeke) 50ml
100U/ml penicillin, 0.1mg/ml streptomycin (Sigma-Aldrich®, Bornem) 5ml
2. 2D culture
VEC and VIC were cultivated in a monolayer culture in an incubator (37°C, 5% CO2). Every 3
days medium was refreshed. When confluency was reached, cells were passaged using
0.025% Trypsin (Gibco®) and 0.01% EDTA-solution at 37°C for approximately 10min after
washing the cells twice with PBS. Cells from passage 3 to 6 were used for all the experiments.
3. 3D culture
3.1. Agarose microwell chips
To synthesize the microwell chips, 3 w/v% agarose powder (UltrapureTM Agarose; Life
TechnologiesTM) was dissolved by heating in PBS. The agarose solution was then poured into
polydimethylsiloxane (PDMS) molds in a 6-well plate and centrifuged at 3000rpm for 60s (Fig.
9). After cooling at room temperature, the molds were manually removed. The chips were then
15
punched into smaller size and transferred to a 12-well plate. The produced chips were
incubated in PBS and stored in the refrigerator.
3.2. Micro-aggregate formation
Micro-aggregates were generated by seeding single cells in agarose microwells. Prior to the
seeding, PBS in the microwells was replaced by 0.5ml cDMEM+ supplemented with 250µM
ascorbic acid 2-phosphate (cDMEM++ ASC-2P) medium and pre-warmed at 37°C. Confluent
VIC cultures were harvested (see section 2.). Afterwards, the number of cells was determined
by using a Bürker counting chamber (Marienfeld Superior). To distinguish between viable and
dead cells, cells were first stained with trypan blue (Gibco®) before cell counting. Cells were
resuspended in cDMEM++ASC-2P and predefined number of cells (1*106, 1.5*106, 2*106) were
gently seeded (Fig. 8) in each microwell (pore size 400µm), followed by adding 2-3ml extra
medium after at least 1h. The empty wells were filled with PBS and the seeded microwells
were maintained in an incubator at 37°C and 5% CO2. Medium was refreshed after 1 day, then
every 2 days.
Figure 8: Schematic overview of agarose microwell fabrication and micro-aggregate generation
3.3. Micro-aggregate collection
VIC micro-aggregates were collected by powerful pipetting/resuspending of the medium. This
method was repeated for several times until the microwells were empty. Microwells were then
rinsed 2 times with PBS. The collected micro-aggregates were pooled in a 50ml tube and
centrifuged at 1000rpm for 5min. Afterwards, the supernatants was removed and the micro-
aggregates pellet was used for experiments. Aggregate diameter (n=20) was determined with
the Xcellence image software.
3.4. Methacrylamide-modified gelatin (Gel-MOD) synthesis
Gelatin methacrylamide was prepared by reaction of Bovine type B gelatin (Rousselot) with
methacrylic anhydride. After dissolution of gelatin in phosphate buffer (pH 7.8) at 50 °C,
16
methacrylic anhydride was added while vigorously stirring. After 1 h of reaction, the reaction
mixture was diluted and dialyzed for 24 h against distilled water at 40 °C. Purified gelatin
methacrylamide had a degree of substitution of 61.72%, as determined by 1H NMR (Bruker
AVANCE II 500 MHz). Prior to use, the freeze-dried Gel-MOD building blocks were sterilized
by ethylene oxide treatment (cold cycle, Maria Middelares hospital, Ghent).
3.5. Cell/micro-aggregate encapsulation in Gel-MOD
Single cells and micro-aggregates were encapsulated in methacrylamide-modified gelatin Gel-
MOD hydrogels. Two concentrations of Gel-MOD (10 w/v% and 20 w/v%) and two PI (2mol%
Irgacure 2959 and 20mol% VA-086) were used in this study (Table 2).
Table 2. Various conditions of cell/microaggregates-laden Gel-MOD
w/v% gels + PI
Gel-MOD (mg/ml PBS)
Amount PI (µl)
Density single cells
Density micro-aggregates
Gel-MOD solution/gel
PG (min)
CL (min)
10% + Irg 100 13.3 5.106/ml Min. 6 chips/ml 250µl 10 20
20% + Irg 200 26.6 5.106/ml Min. 6 chips/ml 250µl 10 20
10% + VA 100 34.2 5.106/ml Min. 6 chips/ml 250µl 10 20
20% + VA 200 68.4 5.106/ml Min. 6 chips/ml 250µl 10 20
Gel-MOD was dissolved overnight in PBS on a shaker at 37°C, and the PI was prepared on
the day of the experiments. 0.8 w/v% of Irgacure and 4 w/v% of VA-086 were dissolved in PBS
respectively at 65°C and in 37°C for approximately 1h or until totally dissolved. PI was sterilized
by a 0.22µm filter (Milipore) before it was added to the Gel-MOD solution.The amount of the
added PI was calculated based on the formula on Table 3.
Table 3. Formula of PI amount
Irgacure 2959 VA-086
0.000385 mol amine groups per gram Gel-MOD * 0.6172 (DS) * 0.02 mol (used Irgacure concentration) * 224.3g/mol (MW Irgacure) ÷ 0.008g/ml (w/v of Irgacure) = 133µl per gram Gel-MOD
0.000385 mol amine groups per gram Gel-MOD * 0.6172 (DS) * 0.2 mol (used VA-086 concentration) * 288.35g/mol (MW VA-086) ÷ 0.04g/ml (w/v of VA-086) = 342µl per gram Gel-MOD
Shortly before the encapsulation, 2 mol% Irgacure 2959 (0.01 and 0.02 w/v%, respectively for
10 and 20 w/v% Gel-MOD) or 20 mol% VA-086 (0.13 and 0.26 w/v%, respectively for 10 and
20 w/v% Gel-MOD) was added to the Gel-MOD solution and vortexed. This solution was then
mixed with single cells or micro-aggregates and gels were formed in a 48 well plate (250µl in
each well) (Table 2). Physical gelation (PG) was allowed at the room temperature for 10min,
followed by chemical crosslinking (CL) of the methacrylate groups with 365nm UV-A light
(4mW/cm2, UVP Inc.) for 20min. After crosslinking, cDMEM++ASC-2P was added to the
17
hydrogels. The single cells/micro-aggregates-laden hydrogels were replaced in 6-well plates
to provide more supply of medium and the medium was refreshed every 3 days.
3.6. Fusion of micro-aggregates in Gel-MOD
Based on the method of encapsulation (section 3.5.), 10 w/v% Gel-MOD was chosen to
analyze the potential of micro-aggregates fusion in Gel-MOD. In these experiments, micro-
aggregates from 200 chips/ml were encapsulated in 250µl 10%w/v Gel-MOD and crosslinked
with UV irradiation in the presence of VA-086 and gels were formed in a 48-well plate (Table
4).
Table 4. Conditions in fusion experiments
%w/v gels + PI Density micro-
aggregates Gel-MOD
solution/gel PG (min) CL (min) Well-plate
10% + VA Min. 200 chips/ml 250µl 5 at 4°C 20 48
4. Viability assay
The viability of single cells and micro-aggregates (in Gel-MOD) were determined by using
Calceine-AM (AnaSpec; 89201)/Propidium Iodide (Sigma-Aldrich®; P4170) (Ca/Pi) staining.
Living cells were stained green, while dead cells were stained red. Before the application, the
single cells/micro-aggregates-laden hydrogels or micro-aggregates (from 2 chips) were rinsed
with PBS. For this staining, 1ml Ca/Pi solution (2μl Pi and 2μl calcein-AM in every ml PBS)
was added to the single cells/micro-aggregates-laden hydrogels or micro-aggregates and the
samples were incubated in the dark for 10min at room temperature. The results were analyzed
by a fluorescence microscope (Olympus IX 81).
5. Histochemical staining
Micro-aggregates-laden hydrogels and micro-aggregates (from 2 chips) were fixed in 4%
paraformaldehyde overnight at 4°C. Afterwards, the samples were dehydrated sequentially in
a set of increasing alcohol concentration (30%, 50%, 70%, 96% disolol + 1% toluidine blue,
isopropyl alcohol, isopropyl alcohol/toluol) for 1h and toluol for 30min, embedded in paraffin
(Thermo ScientificTM), and placed overnight in an oven (60°C). 5µm-thick sections were cut
from the samples for (immuno)histochemical stainings. Prior to staining, samples were
deparaffinized and after staining, samples were dehydrated sequentially in a set of increasing
alcohol concentration and toluol. The last step was mounting a coverslip on the glass slide with
mounting solution (Richard Allan Scientific).
5.1. Hematoxylin-eosin staining
Hematoxylin-eosin (HE) staining is a principal histological staining that allows to visualize the
tissue morphology. This staining uses two dyes, hematoxylin (basic dye) which gives purplish
blue color to the cell nuclei, and eosin (acidic dye) which stains cytoplasm into reddish or pink.
18
The HE staining was performed by a robot-stainer (HMS740 MICROM robot-stainer, Walldorf,
Duitsland) (Programme: 3x5min toluol; 2x2min isopropyl alcohol; 2x2min 96% alcohol; 2min
tap water; 1min distilled water; 15s haematoxyline (VWR®); 2min tap water; 1min clarifier I
(Thermo-scientificTM); 1min tap water; 1min bluing reagent (Thermo-scientificTM); 1min tap
water; 1min distilled water; 30s 1% eosine+phloxine (Thermo-scientificTM); 2min tap water;
2x2min 96% alcohol; 2x2min isopropyl alcohol; 2x2min toluol; 1min toluol).
5.1.1. Alcian blue (GAG)
Alcian blue (AB) was used to detect GAG. For this staining, the samples were incubated in 3%
acetic acid (Chem-lab) AB (VWR international) solution for 30min at the room temperature and
then rinsed by distilled water.
5.1.2. Silver staining (Reticular fibers)
Gordon and Sweet’s silver staining was applied to visualize the reticular fibers. This staining
was done by incubating the samples sequentially for certain durations in several kinds of
solution that were freshly prepared. The first step was oxidizing the samples in acidified
potassium permanganate for 8min. Afterwards, the samples were rinsed under streaming tap
water. The samples were then decolorized in 1% oxalic solution for 1min and rinsed again
under streaming tap water. This step is followed by rinsing the samples 2x1min in distilled
water. Next, sample were incubated in 2.5% Iron alum for 25min, rinsed 3 times in distilled
water, and incubated in ammoniacal silver nitrate solution for 30s. Subsequently, samples
were shortly rinsed in aquadest and incubated in 4% formadehyde for 30s, then rinsed again
in distilled water for 1m. The following step was incubating the samples in 0.2% gold chloride
for 3min and rinsed in distilled water for 1min. The samples were fixed in 5% sodium
thiosulphate for 5min and rinsed under streaming tap water for 15min. Finally, samples were
counterstained by incubation in nuclear fast red for 8min and rinsed in distilled water. Nuclei
stained red and reticular fibers black.
5.2. Immunohistochemical staining
Immunohistochemical staining (IHS) was performed to demonstrate particular antigens that
are present in the samples by using specific antibodies. In this study, indirect IHS was used.
In other words, the primary antibodies bind on the antigens and the labelled-secondary
antibodies will recognize the primary antibodies. IHS against Col I and Ki67 were performed
to detect the presence of collagen I fibers and cell proliferation.
Prior to the IHS, citrate buffer pH 6.0 was prepared and cooked for 5min in a microwave. The
deparaffinized samples were cooked in the heated citrate-buffer for 2x5min in a microwave.
After cooling for 20min, the samples were lined with toluol-marker (Dako) and washed 2x5min
with PBS. Meanwhile, blocking serum (BS) that contained PBS, 5% normal rabbit serum (NRS;
Dako), 1% bovine serum albumin (BSA; Roche Diagnostics, Vilvoorde, Belgium), and 0.2%
19
Tween20 (VWR®) was prepared and diluting buffer (VB) was made by mixing BS and PBS in
proportion 1:10.
The samples were pre-treated with 3% H2O2 ((VWR®) for 10min. Afterwards, the samples
were treated with BS for 30min, followed by incubating them with the collagen type I primary
antibody (anti-mouse monoclonal; Santa Cruz sc-293182; 1/50 diluted in VB) or Ki67 primary
antibody (anti-mouse monoclonal; Dako M7240; 1/50 diluted in VB). For Col I staining, the
samples were incubated with the primary antibody and kept overnight at 4°C, while for the Ki67
staining, samples were incubated with the primary antibody for 2h at room temperature. The
samples were then rinsed 2x5min with PBS. The next step was incubating samples (both
collagen I and Ki67 staining) with the secondary antibody (biotinylated-rabbit-anti-mouse;
Dako; 1/200 diluted in VB) at room temperature. After 30min, the samples were rinsed 2x5min
with PBS and incubated with Streptavidine - Horse Radish Peroxidase (Dako P0397; 1/200
diluted in VB) for 30min at room temperature. The samples were rinsed 2x5min with PBS and
1x5min with Tris-buffer. The last step was incubating the samples with chromogenic substrate
that contained Tris Buffer, 0,06% Diaminobenzidine (DAB) (Sigma-Aldrich®), and 0,03% H2O2
(VWR® ) for 10min at room temperature in the dark. The reaction of chromogenic substrate
was stopped by putting the samples under streaming tap water for 10min. The samples were
counterstained with 25% Mayer’s hematoxylin.
6. RT-qPCR
For RT-qPCR, RNA from VIC monolayer (6*106 cells), VIC aggregates (from 5-6 chips
containing 1*106cells/chips), and aggregates-laden Gel-MOD (from 10-12 chips). Prior to RNA
isolation, Gel-MOD was degraded by collagenase type IA (Sigma-Aldrich®; ≥125U/mg) at
37°C for approximately 1h or until the gel was dissolved. Each sample was lysed in 350µl RLT
buffer and stored at -20°C. The purification of the isolated RNA was performed by using
RNeasy mini kit (Qiagen). The purity of the isolated RNA was measured with a NanoDrop ND-
1000 spectrofotometer (Thermo ScientificTM) and the RNA concentration was diluted to
50ng/µl, and then stored at -80°C. The next step was converting the RNA into cDNA. 10μl
2xMastermix was added to 10µl RNA. The mastermix consisted of 10x RT buffer, 25x dNTP
Mix, 10x RT random primers, MultiScribe Reverse transcriptase, RNase inhibitor en nuclease-
free water (High Capacity cDNA Reverse Transcription Kit, Life Technologies). The prepared
samples were then inserted in the PCR machine to run the reverse transcription (Program:
25°C,10min; 37°C,2h; 85°C,5min; 4°C,∞). The generated cDNA could be stored at -20°C.
SYBR select Master Mix (Life Technologies) was used to amplify the cDNA and to perform the
qPCR reaction. In each well of 96well-optical reaction plate 10µl SYBR Select Master Mix,
0.6µl forward primer, 0.6 reverse primer, 6.8µl RNAse-free water, and 2µl cDNA template were
20
added. Prior to the reaction, the well-plate was centrifuged at 1000rpm for 30s. The reaction
was performed in the PCR machine with following program settings: 50°C, 2min; 95°C, 2min;
40 cycles of 95°C, 15s; 60°C, 1 min. Table 5 presents the forward and reverse primer
sequences.
Table 5. Forward and reverse primer sequences for porcine valvular interstitial cells (VIC)
Gene Symbol
Forward Primer 5’ 3’
Reverse Primer 3’ 5’
Amplicon Length
(bp)
Primer Efficiency
(%) EEF1A1 CAACATGCTGGAGCCAAGTG AGTGTGGTTCCACTGGCATT 87 94.2
HPRT1 GGGAGGCCATCACATCGTAG CGCCCGTTGACTGGTCATTA 167 89.6
α-SMA GGACCCTGTGAAGCACCAG GTCACCCACGTAGCTGTCTT 198 98.9
VIM TCTGGAATCCCTCCCTCTGG TCGCTGCACAGAGTACATGC 146 93.8
FSP-1 TACTCAGGCAAGGAGGGTGA TGGAAAGCAGCTTCATCCGT 113 94.3
OCN TACCCAGATCCTCTGGAGCC TGCCATAGAAGCGCCGATAG 109 110.4
Egr-1 CTATCAAGGCCTTTGCCACG AGCTCATCTGAGCGGGAGA 193 104.4
Sox-9 CATCTCTCCCAACGCCATCT TCTCGCTTCAGGTCAGCCTT 174 94.6
Runx2 CAACTTCCTGTGCTCTGTGCT GAGAACCAGGGTTGAGGTGAT 90 93.7
COL1A1 AGACATCCCACCAGTCACCT TCACGTCATCGCACAACACA 122 96.9
COL3A1 GCTCCCATCTTGGTCAGTCC ATGGGATCTCAGGGTTGGGA 147 94.5
COL5A1 GCTCAGCGTCCACAAGAAGA GTGGTCCGAGACAAAGAGCA 184 91.4
Elastin CCCACACATCGAAACCCTCA CACAAGAGCAAGGGGGCTAA 155 94
HA GTGATTGCCTCCTTCGACCA TGCACAGACCCATCACTGAG 89 94.4
Biglycan TCACTGGCATCCCCAAAGAC AGGGTGGGCAGAAAACTCAG 178 97.4
Decorin CCTGCAAACTCTTGCTTGGG TAATCCGGGGATTTGCCACA 81 94.6
Versican AACCAGACAGGCTTCCCTTC AAGTGGGTGAGGCAGTTTCT 124 92.7
MMP-1 GGCCATCTATGGACCTTCCG TGTGCGCATGTAGAACCTGT 145 101.5
MMP-2 CGATGGCTTCCTCTGGTGTT AGCTGTTGTAGGATGTGCCC 155 94
MMP-3 CACTTACAGACCTGGCTCGG AGATTCTGTGGGCTCAACGG 113 89.3
MMP-9 ACTTCGGAAACGCAAAAGGC AAGAGTCTCTCGCTAGGGCA 169 110.9
TIMP-1 CACCTGCAGTTTTGTGGCTC GGGATGGATGTGCAGGGAAA 117 89.5
TIMP-2 CTCCGGGAACGACATCTACG GCCTTTCCTGCGATGAGGTA 180 93
7. Gel-MOD characterization
7.1.1. Equilibrium swelling test
This test was performed to analyze the effect of PI (2 mol% Irgacure 2959 or 20 mol% VA-
086) and Gel-MOD concentration (10 and 20 w/v%) on the crosslinking and swelling behavior
of gels. Hydrogels disks (Ø = 15mm; height = 1-2mm) were prepared by pipetting 300ml Gel-
MOD solution in each well of a 24-well plate and crosslinked for 20min with 365nm UV-A light
(4mW/cm2, UVP Inc.). After crosslinking, the disks were freeze-dried at 0.370 atm and -50°C
in a Christ Alpha1-4 freeze-dryer and weighed (Wd0). The dried disks were then immersed in
DMEM at 37°C until equilibrium swelling was reached. At this temperature, non-covalent
crosslinked polymer chains will leach out from the hydrogels by diffusion. Consequently, disks
were removed from the medium and gently wiped with paper, and weighted again (Whe).
Subsequent lyophilization gave the final dry weight (Wde). The gel fraction is calculated as the
percentage of the hydrogel that is covalently incorporated into the 3D network according to
equation 1 (Eq. 1), while the equilibrium mass swelling ratio (%) is calculated according to Eq.
2 [66, 67]. All measurements were performed in triplicate.
21
%𝐺𝑒𝑙 𝑓𝑟𝑎𝑐𝑡𝑖𝑜𝑛 = 𝑊𝑑𝑒
𝑊𝑑0. 100 (Eq. 1)
%𝑆𝑤𝑒𝑙𝑙𝑖𝑛𝑔 =𝑊ℎ𝑒− 𝑊𝑑0
𝑊𝑑0. 100 (Eq. 2)
7.1.2. Compression test
Crosslinked hydrogel disks (Ø = 15mm; height = 1mm) were compressed at room temperature
using a TA500 texture analyzer (Lloyd Instruments) equipped with a 10N load cell. The
compression was applied over a distance of 0.5mm for 10 cycles at a rate of 5mm/min.
Afterwards, the compressive modulus was determined from the slope of the stress-strain curve
during the first cycle. All measurements were performed in triplicate.
8. Data analysis
The normality of the variables was determined by the Saphiro-Wilk test. All variables were
parametric data and the statistical analysis was performed using a one-way analysis of
variance (ANOVA) with a post-hoc Tukey’s using the commercially available software package
SPSS for windows, version 23.0 (SPSS GmbH Software, München, Germany). Statistically
significant data were characterized with *p<0.05 or ***p<0.001 in comparison with the standard
monolayer cultured VIC, and #p<0.05 or ###p<0.001 for significant differences between groups.
22
RESULTS
1. Encapsulated single cells in Gel-MOD
VIC single cells were encapsulated in 10 and 20 w/v% Gel-MOD and crosslinked with UV
irradiation in the presence of 2 mol% Irgacure 2959 or 20 mol% VA-086 PI. The encapsulated
VIC were incubated in medium with or without ASC-2P. The viability of VIC single cells in the
gels and on the surface of the gel was evaluated by live/dead assay (Ca/Pi staining) at day 1,
3, and 7 (Fig. 9).
Figure 9: Live/dead staining of encapsulated VIC single cells in Gel-MOD (in the gel (IN) and on the surface of the gel (ON)) at day 1, 3, and 7. (A) 10 w/v% Gel-MOD + Irgacure 2959 (B) 20 w/v% Gel-MOD + Irgacure 2959 (C) 10 w/v% Gel-MOD + VA-086 (D) 20 w/v% Gel-MOD + VA-086 [Scale bar = 200µm, magnification 10x]
In general, VIC single cells on the surface of the gels were elongated, while VIC in the gels
were circular. After 7 days, there were no or few viable cells in the gels in all conditions. In
contrast, cells on the surface of the gels were still viable and highly proliferative. VIC in gels
that were cultivated in the presence of ASC, showed higher viability at the earliest time points
(day 1 and 3), however, after 7 days also no or few viable cells were observed (Fig. 9A-D).
Fig. 9 also showed that cells on the surface of gels were more proliferative on 10 w/v%
compared to 20 w/v% gels and extremely more proliferative on VA-086 crosslinked gels.
23
2. VIC micro-aggregates
The generation of VIC micro-aggregates had already been optimized in the preliminary study
(master thesis Inès Puype), where VIC micro-aggregates were generated from 1*106 cells/chip
with Ø400µm pores and cultured in cDMEM+ with 250µM ASC-2P. In this study, VIC micro-
aggregates were generated from 1*106, 1.5*106, and 2*106 cells/chip with Ø400µm pores. The
morphology, viability, proliferation, and ECM production was studied at day 3, 7, 14, and 21
(Supplement S1), while the comparison between day 3 and day 21 was depicted in Fig. 10.
In addition, Table 6 provides the mean number of cells/micro-aggregates generated from
1*106, 1.5*106, and 2*106 cells/chip with Ø400µm pores.
Table 6. Number of cells/micro-aggregate [Adapted from [68])
Seeded number of cells/chip
Mean number of cells/micro-aggregate
1*106 629
1.5*106 944
2*106 1258
The morphology of VIC micro-aggregates was observed by a phase-contrast microscope and
HE staining. As shown in Fig. 10A, VIC micro-aggregates from the three different conditions
all shrank in size by day 21. This result was confirmed by the size measurement of the micro-
aggregates, which showed that the diameter of micro-aggregates from 1*106 cells/chip
decreased from 150µm at day 3 to 110µm at day 21, while the diameter of micro-aggregates
from 1.5*106 and 2*106 cells/chip decreased from and ±200µm to ±140µm (Fig. 11). There
was no remarkable difference of micro-aggregates diameter from 1.5*106 and 2*106 cells/chip
during the study, while the diameter of micro-aggregates from 1*106 cells/chip was strikingly
lower than the other conditions at all time points (Fig. 11). Furthermore, VIC micro-aggregates
had a clear contour, circular form, and no pyknotic nuclei as seen from HE staining from day 3
up to day 21 (Fig. 10C and S1A).
Live/dead (Ca/Pi) staining was performed on VIC micro-aggregates to study the viability of the
micro-aggregates. In all conditions, VIC micro-aggregates were viable up to 21 days (Fig. 10B
and S1B). The proliferation of VIC micro-aggregates was analyzed by Ki67 immuno-staining
(Fig. 10D). VIC micro-aggregates in all conditions were sporadically positive for Ki67. The
Ki67+ cells were specifically located at the outer edge of the aggregates. At day 3, VIC micro-
aggregates generated from 1*106 cells/chip showed more Ki67+ cells compared to the bigger
aggregates. However, after 7 days the number of Ki67+ cells dropped with time in all conditions
(Fig. 10D and S1B).
24
Figure 10: Morphology (A: bright field (BF) microscopy and C: HE staining), viability (B: Ca/Pi staining), proliferation (D: Ki67 immuno-staining), and ECM production (E: Col I staining; F: AB staining (GAG); G: Silver staining (reticular fibers)), of VIC micro-aggregates from 1*106, 1.5*106, and 2*106 cells/chip at day 3 and day 21 [Scale bar = 200µm, magnification 10x (A and B)], [Scale bar = 40µm, magnification 40x (C-G)]
The ECM production (Col I, GAG, and reticular fibers) of VIC micro-aggregates is illustrated in
Fig. 10E-G and S1C. At day 21 all aggregates showed Col I fibrils and these fibrils were most
apparent in aggregates from 1*106 cells/chip. Col I staining was not visible at day 3 but started
to increase from day 14. From this time point on, small fibrils became visible and more
pronounced at the outer edges of the aggregates (Fig. 10E and S1C). AB staining showed
that GAG production in aggregates increased with the study duration in all conditions. The
highest GAG content on day 3 was found in the center of VIC micro-aggregates from 1*106
cells/chip, by day 21 GAG were present throughout the whole micro-aggregates. On the other
hand, micro-aggregates from 1.5*106 and 2*106 cells/chip, showed less GAG at day 3 and
GAG production was more apparent at the periphery of micro-aggregates as shown at day 14
25
and day 21 (Fig. 10F and S1C). Reticular fibers were abundantly present as early as day 3
throughout the whole body of the micro-aggregates in all conditions and started to organize in
circular bundles towards the edges from day 7 on (Fig. 10G and S1C).
Figure 11: VIC micro-aggregates diameter from 1*106, 1.5*106, and 2*106 cells/chip from day 3 until day 21, n=20. Data is presented as mean and error bars indicate the 95% confidence interval.
3. Encapsulated VIC micro-aggregates
VIC micro-aggregates from 1*106 cells/chip were encapsulated in 10 and 20 w/v% Gel-MOD
and crosslinked with UV light irradiation in the presence of Irgacure 2959 or VA-086 PI. The
morphology, viability, proliferation, and ECM production of the encapsulated micro-aggregates
was evaluated at day 1, 3, 7, and 14.
VA-crosslinked gels were found to be less stable than Irgacure-crosslinked gels. 10 w/v% VA-
crosslinked gels started to disintegrate at day 7 and 20 w/v% VA-crosslinked gels started to
disintegrate at day 14, whereas 10 w/v% Irgacure-crosslinked gels started to disintegrate at
day 14 and 20 w/v% Irgacure-crosslinked gels were still stable until day 21 (data not shown).
The encapsulated micro-aggregates were still viable up to 7 or 14 days in the VA-crosslinked
gels (Fig. 12B&F) as well as in Irgacure-crosslinked gels (Fig. 14B&F). Furthermore, HE
staining showed no pycnotic nuclei in the micro-aggregates (Fig. 12C&G and Fig. 14C&G).
Nevertheless, it is noteworthy that VIC micro-aggregates in 20 w/v% Irgacure-crosslinked gels
seemed to have more matrix proportion and less nuclei at day 14 (Fig. 14G). Micro-aggregates
150
136
125
110
184
165
153
140
198
171
160
144
100
120
140
160
180
200
220
D3 D7 D14 D21
Dia
me
ter
(µm
)
Duration (days)
Size of VIC micro-aggregates
1 million cells/chip 1.5 million cells/chip 2 million cells/chip
26
maintained their circularity in the gels (Fig. 12E&G and Fig. 14A, C, E, G), except in 10 w/v%
VA-crosslinked gels where cells from the encapsulated micro-aggregates started to migrate in
the gel in all directions (Fig. 12A&C). Ki67+ cells were found at the periphery of the micro-
aggregates at day 1 and 3, and also in migrating cells by day 7 (Fig. 12D). Ki67+ cells were
also found at the outer edges of VIC micro-aggregates in other gels at day 1, but the amount
of positive cells rapidly decreased over time (Fig. 12H and Fig. 14D&H).
Fig. 13 and Fig. 15 depicted the ECM production of VIC micro-aggregates respectively in VA-
crosslinked gels and in Irgacure-crosslinked gels. The encapsulated micro-aggregates were
negative for Col I in all conditions during the study (Fig. 13A&D and Fig. 15A&D). Furthermore,
at day 1 only a limited amount of GAG was visualized with AB staining, however, from day 3
there was a rapid time dependent increase in GAG staining throughout the micro-aggregates.
Compared to micro-aggregates in 10 w/v% gels, micro-aggregates in 20 w/v% showed more
intense AB staining (Fig. 13B&E and Fig. 15B&E). On the other hand, reticular fibers were
abundantly present from day 1 in all conditions and started to align over time towards the outer
edge of the micro-aggregates. It is also noteworthy that micro-aggregates which located close
to each other appeared to fuse and that the reticular fibers of these aggregates started to move
towards each other (except in 20 w/v% Irgacure-crosslinked gels) (Fig. 13C&F and Fig.
15C&F).
27
Figure 12: Morphology (A&E: bright field (BF) microscopy and C&G: HE staining), viability (B&F: Ca/Pi staining), and proliferation (D&H: Ki67 immuno-staining) of encapsulated VIC micro-aggregates in 10 and 20 w/v% VA-crosslinked Gel-MOD at day 1, 3, 7, and 14 [Scale bar = 200µm, magnification 10x (A, B, E, F)], [Scale bar = 40µm, magnification 40x (C, D, G, H)]
28
Figure 13: ECM production (A&D: Col I staining; B&E: AB staining (GAG); C&F: Silver staining (reticular fibers)) of encapsulated VIC micro-aggregates in 10 and 20 w/v% VA-crosslinked Gel-MOD at day 1, 3, 7, and 14 [Scale bar = 40µm, magnification 40x]
29
Figure 14: Morphology (A&E: bright field (BF) microscopy and C&G: HE staining), viability (B&F: Ca/Pi staining), and proliferation (D&H: Ki67 immuno-staining) of encapsulated VIC micro-aggregates in 10 and 20 w/v% Irgacure-crosslinked Gel-MOD at day 1, 3, 7, and 14 [Scale bar = 200µm, magnification 10x (A, B, E, F)], [Scale bar = 40µm, magnification 40x (C, D, G, H)]
30
Figure 15: ECM production (A&D: Col I staining; B&E: AB staining (GAG); C&F: Silver staining (reticular fibers)) of encapsulated VIC micro-aggregates in 10 and 20 w/v% Irgacure-crosslinked Gel-MOD at day 1, 3, 7, and 14 [Scale bar = 40µm, magnification 40x]
31
4. Gene expression of encapsulated VIC micro-aggregates
After three days of micro-aggregates culture (from 1*106 cells/chip), aggregates were
encapsulated in 20 w/v% Irgacure-crosslinked Gel-MOD. RT-qPCR was performed to analyze
the mRNA expression levels of ECM components (Col I/III/V, elastin, HA, biglycan, decorin),
MMPs/TIMPs (MMP-1/2/3/9, TIMP-1/2), and phenotype markers (α-SMA, vim, fibroblast
specific protein-1 (FSP-1), osteocalcin, Egr-1, Sox-9, Runx2) in VIC micro-aggregates in gels
at an early and a late time point, respectively day 3 and 14.
To compare the expression levels of encapsulated aggregates with the expression levels of
the target genes in non-encapsulated micro-aggregates at the same time points, RNA was
isolated from the non-encapsulated micro-aggregates at day 7 and 18. The expression levels
of the target genes were normalized to the geometric mean of two reference genes, Eukaryotic
translation elongation factor 1 alpha 1 (EEF1A1) and Hypoxanthine phosphoribosyltransferase
1 (HPRT1). The relative gene fold changes were determined by the 2ΔΔCt method (compared
to the mRNA expression level of standard monolayer cultured VIC) to determine the influence
of the 3D culture and the gels on VIC mRNA expression.
Fig. 16 shows that the mRNA expression of the main ECM components produced by VIC,
increased in non-encapsulated as well as in encapsulated VIC micro-aggregates when
compared to standard monolayer cultured VIC (significant differences are marked with an
asterix). Col I, III, and V mRNA expression levels increased over time in the encapsulated
micro-aggregates. However, the increase was only significant for Col I and V. On the other
hand, the mRNA expression levels of Col I, III, and V in the non-encapsulated micro-
aggregates remained unchanged over time. Time-dependent increase of elastin, non-sulfated
GAG (HA), and PG (versican, biglycan, and decorin) expression was positively observed in
the encapsulated micro-aggregates. Moreover, the expression levels of these ECM
components were higher compared to the non-encapsulated micro-aggregates. Nevertheless,
this was only significant for biglycan. The non-encapsulated micro-aggregates, on the other
hand, only showed a time-dependent increase of HA, versican and biglycan expression levels.
MMP-1/2/3/9 and TIMP-2 expression in the non-encapsulated increased over time. The
encapsulated micro-aggregates, on the other hand, showed only a time-dependent increase
of mRNA expression levels for MMP-1/3/9, while MMP-2 and TIMP-2 remained stable. It is
noteworthy that the non-encapsulated micro-aggregates showed a more pronounced increase
of MMP-1/2/3/9 and TIMP-2 expression compared to the encapsulated micro-aggregates and
these expression levels were significantly higher in comparison with the standard monolayer
cultured VIC. In contrast to the expression of TIMP-2, the non-encapsulated micro-aggregates
exhibited a significant decrease of TIMP-1 over time, while no significant difference of TIMP-1
expression was observed in the encapsulated micro-aggregates (Fig.17).
32
Figure 16: Relative mRNA expression level of ECM components in VIC micro-aggregates at day 7 and 18, as well as encapsulated VIC micro-aggregates in 20 w/v% Irgacure-crosslinked Gel-MOD 2959 at day 3 and 14 compared to monolayer cultured VIC, n=3. Data is presented as mean and error bars indicate the 95% confidence interval. Statistically significant data are marked with *p<0.05 or ***p<0.001 in comparison with the standard cultured monolayer, and #p<0.05 or ###p<0.001 for significant differences between groups.
Figure 17: Relative mRNA expression level of MMPs and TIMPs in VIC micro-aggregates at day 7 and 18, as well as encapsulated VIC micro-aggregates in 20 w/v% Irgacure-crosslinked Gel-MOD at day 3 and 14 compared to monolayer cultured VIC, n=3. Data is presented as mean and error bars indicate the 95% confidence interval. Statistically significant data are marked with *p<0.05 or ***p<0.001 in comparison with the standard cultured monolayer, and #p<0.05 or ###p<0.001 for significant differences between groups.
33
Fig. 18 demonstrates the expression level of several markers used to study the phenotype of
the non-encapsulated and the encapsulated micro-aggregates. α-SMA expression in all
conditions was significantly lower than α-SMA expression of the standard monolayer cultured
VIC, while vim expression levels were slightly, but not significant, higher than the standard
monolayer cultured VIC, and no significant difference was observed between the non-
encapsulated and the encapsulated micro-aggregates over time. FSP-1 and OCN expression
levels were higher in micro-aggregates compared to the standard monolayer cultured VIC, and
expression levels remained stable over time. Egr-1, Sox-9, and Runx2 showed various
expression levels in the non-encapsulated as well as in the encapsulated micro-aggregates.
Egr-1 expression levels in the encapsulated micro-aggregates increased over time, while
decreased in the non-encapsulated micro-aggregates. Sox-9 was significantly more expressed
in the encapsulated micro-aggregates compared to the non-encapsulated micro-aggregates
and the standard monolayer cultured VIC and remained relatively unchanged over time. On
the other hand, Runx2 expression increased in the non-encapsulated micro-aggregates and
was significantly higher than in the encapsulated micro-aggregates and the standard
monolayer cultured VIC.
Figure 18: Relative mRNA expression level of markers in VIC micro-aggregates at day 7 and 18, as well as encapsulated VIC micro-aggregates in 20 w/v% Irgacure-crosslinked Gel-MOD at day 3 and 14 compared to monolayer cultured VIC, n=3. Data is presented as mean and error bars indicate the 95% confidence interval. Statistically significant data are marked with *p<0.05 or ***p<0.001 in comparison with the standard cultured monolayer, and #p<0.05 or ###p<0.001 for significant differences between groups.
34
5. Fusion of VIC micro-aggregates in Gel-MOD
Micro-aggregates can be used as building blocks to form a larger tissue construct by self-
assembly. Based on the results mentioned above, a proof of concept fusion experiment is
performed to study the capability of the encapsulated micro-aggregates to form larger valvular
micro-tissues. 10 w/v% VA-crosslinked Gel-MOD was chosen to encapsulate a huge number
of VIC micro-aggregates from 1*106 cells/chip. In this experiment, observations on the
morphology, fusion, proliferation, migration, and ECM production of the encapsulated micro-
aggregates were performed at day 3 and 7.
By day 3 VIC micro-aggregates had fused already to form larger valvular micro-tissues in some
regions of the gels. The valvular micro-tissues had normal morphology with round nuclei. After
7 days, some micro-tissues showed pycnotic nuclei and the micro-aggregates that did not fuse
had more ECM proportion and less nuclei. Valvular micro-tissues at day 3 showed more Ki67+
cells in comparison with the micro-tissues at day 7. Similar to the encapsulated micro-
aggregates, the micro-tissues were not positive for Col I at day 3 and 7, showed a time-
dependent increase of GAG, and produced massive reticular fibers at early time point. In
addition, GAG were detected throughout the micro-tissues and reticular fibers were randomly
aligned throughout the micro-tissues (Fig. 19).
35
Figure 19: Morphology (HE staining) and ECM production (Col I staining, AB staining (GAG), and Silver staining (reticular fibers)) of encapsulated VIC micro-aggregates in 10 w/v% VA-crosslinked Gel-MOD at day 3 and 7 [Scale bar = 100µm, magnification 10x (A&C)], [Scale bar = 200µm, magnification 20x (B&D)]
6. Gel-MOD characterization
An equilibrium swelling test was performed to analyze the effect of PI (2 mol% Irgacure 2959
or 20 mol% VA-086) and Gel-MOD concentration (10 and 20 w/v%) on the crosslinking and
swelling behavior of gels. Fig. 20 shows that Gel-MOD in all conditions excessively swelled
from the initial state after being immersed in medium. The swelling ratio significantly increased
with decreasing Gel-MOD concentration and VA-crosslinked gels showed significantly higher
(581.5-1081.2%) swelling ratio than Irgacure-crosslinked gels (446.9-778.0%). Fig. 21
illustrates the percentage of the hydrogel that is covalently crosslinked and incorporated into
the 3D structure. Irgacure-crosslinked gels had significantly more covalently incorporated
materials (90.5-97.5%) compared to VA-crosslinked gels (85.5-87.3%). In addition, 20 w/v%
Gel-MOD showed higher percentage of gel fraction than 10 w/v% Gel-MOD.
36
Figure 20: Swelling ratio of 10 and 20 w/v% Gel-MOD with 2mol% Irgacure 2959 and 20 mol% VA-086, n=3. Data is presented as mean and error bars indicate the 95% confidence interval. Statistically significant differences between 10 and 20 w/v% Gel-MOD hydrogels are marked with ***p<0.001, and #p<0.05, ###p<0.001 for significant differences between Irgacure- and VA-crosslinked gels.
Figure 21: Gel fraction of 10 and 20 w/v% Gel-MOD with 2mol% Irgacure 2959 and 20 mol% VA-086, n=3. Data is presented as mean and error bars indicate the 95% confidence interval. Statistically significant differences between 10 and 20 w/v% Gel-MOD hydrogels are marked with ***p<0.001, and #p<0.05, ###p<0.001 for significant differences between Irgacure- and VA-crosslinked gels.
The stiffness of the gels is described as compressive modulus (Fig. 22). The stiffness of the
gels was dependent on the concentration of Gel-MOD and the increase of compressive
modulus was significantly more pronounced in Irgacure-crosslinked gels compared to VA-
37
crosslinked gels. In addition, Irgacure-crosslinked gels showed a significantly higher
compressive modulus (31.9-106.6 kPa) than VA-crosslinked gels (5.0-9.6 kPa).
Figure 22: Compressive modulus of 10 and 20 w/v% Gel-MOD with 2mol% Irgacure 2959 and 20 mol% VA-086, n=3. Data is presented as mean and error bars indicate the 95% confidence interval. Statistically significant differences between 10 and 20 w/v% Gel-MOD hydrogels are marked with ***p<0.001, and #p<0.05, ###p<0.001 for significant differences between Irgacure- and VA-crosslinked gels.
38
DISCUSSION
3D bioprinting has the potential to biofabricate a viable tissue construct by directed assembling
of cells and micro-scale building blocks into a larger construct. This study investigated whether
modified gelatin is an appropriate carrier-material for VIC single cells as well as VIC micro-
aggregates, to obtain a bio-ink that can be used for the biofabrication of a larger, macro-scale
valvular tissue.
Over the past ten years, gelatin and hybrid gelatin hydrogels have been investigated for the
encapsulation of VIC single cells [21, 60, 69-74] and used for 3D bioprinting purposes [62].
Gelatin is a denatured product of type I collagen triple helices and has similar properties to
native collagen, such as RGD peptide sequences which are essential to stimulate cell
attachment, cell migration, cell differentiation, and cell growth [60, 75]. However, due to the
thermo-reversible characteristic of gelatin, chemical crosslinking is needed to enhance the
mechanical properties of the 3D construct [52, 60]. For this reason, gelatin can be modified by
incorporating methacrylates group (Gel-MOD) which can be crosslinked with UV irradiation in
the presence of a PI, such as Irgacure 2959 and VA-086. Irgacure has been largely used in
tissue engineering [58] and is considered as the golden standard for Gel-MOD crosslinking
[57]. Irgacure, in its native state, is non-cytotoxic, however, the photogenerated radical form of
Irgacure can decrease cell viability [58]. It has been shown by Rouillard et al. that when the
Irgacure concentration used for the crosslinking of methacrylamide-modified alginate gels was
raised from 0.02 to 0.03 w/v%, the viability of bovine chondrocytes reduced from >60% to
<40% [58]. The optimal concentration of PI to rapidly and efficiently crosslink the gels with a
short irradiation time is absolutely essential to minimize the effect of the radical agents on the
cells [73]. On the other hand, both the inactive and the activated form of VA have shown to be
cytocompatible [58] and to have minimal effect on the cell viability [57]. In the same study,
Rouillard et al. have shown that compared to the Irgacure-crosslinked hydrogels, VA-
crosslinked hydrogels with concentrations up to 0.3 w/v% had >90% viable chondrocytes [58].
In this study, VIC single cells were encapsulated in 10 or 20 w/v% Gel-MOD and
photopolymerization was performed in the presence of 2 mol% Irgacure or 20 mol% VA to
determine the best condition concerning cell viability. Unfortunately, our results showed that
VIC did not remain viable, regardless the concentration of Gel-MOD or the type of PI. This was
unexpected as other studies have already successfully encapsulated VIC in hybrid Gel-MOD
[21, 69-74], but also in pure Gel-MOD hydrogels [21, 60, 76]. The longest follow-up study of
VIC in pure Gel-MOD hydrogels showed 80% cell viability after 7 weeks, which demonstrated
that Gel-MOD is a good encapsulation material [60].
In order to improve cell viability, ASC-2P was added to the culture medium. ASC-2P is a
derivative of ascorbic acid which is able to stimulate cell proliferation and to enhance Col I
39
production of cells [77]. This organic molecule also has an anti-oxidant function which can
protect cells against oxidative stress [78]. In addition, a recently published article from our
research group has shown that ASC-2P is necessary for the generation of viable, high quality
VIC micro-aggregates and that micro-aggregates quickly degenerate and calcify in the
absence of ASC-2P [79]. Nevertheless, the effect of ASC-2P on the cell viability was noticed
only at the earlier time points (day 1 and 3) and after 7 days no remarkable difference was
observed related to the viability of the encapsulated cells cultured in medium with or without
ASC-2P.
In general, the Gel-MOD used in this study did not support cell survival in the gels regardless
the Gel-MOD concentration, the type of PI, or ASC-2P supplementation. There are several
parameters that can influence the encapsulated cell viability: diffusion of nutrients (size of
hydrogel discs and porosity of the gels), crosslinking duration (UV-A irradiation), light intensity
(irradiation dose), PI (type and concentration), cell-cell contacts, sterilization procedures,
material properties (influenced by degree of substitution, concentration of the gels, and PI),
and type of hydrogel. All of these parameters differ between studies which makes it very hard,
almost impossible to make a straightforward comparison between our study and other studies,
and to find a possible explanation as to why the encapsulated VIC in our study did not survive.
We presume that Gel-MOD is very unlikely to be cytotoxic since VIC on the surface of the gels
did survive and proliferate.
Next, we investigated the encapsulation of VIC micro-aggregates. By using micro-aggregates
as building blocks, instead of single cells, a larger scale valvular tissues can be more rapidly
build. Moreover, 3D micro-aggregates more closely resemble the in vivo micro-environment
where cells naturally interact with their neighboring cells and ECM components [44, 45]. 1*106,
1.5*106, and 2*106 cells were seeded in agarose microwell chips and cultivated in medium
supplemented with ASC-2P. This high throughput system rapidly generated high quality VIC
micro-aggregates within 24h. VIC micro-aggregates from 1*106 cells/chip produced GAG and
Col III (reticular fibers), and showed more Col I production and more Ki67+ cells compared to
the micro-aggregates from 1.5*106 and 2*106 cells/chip. For this reason, all micro-aggregate
encapsulation experiments were performed with micro-aggregates from 1*106 cells/chip after
a 3 day aggregates culture to ensure that the micro-aggregates were solid and completely
compacted. Since micro-aggregates-laden Gel-MOD will be used as a bio-ink in future
bioprinting experiments, it was also important to take the diameter of the micro-aggregates into
account. Micro-aggregates from 1*106 cells/chip had a 150µm diameter at day 3 and these
micro-aggregates can easily pass through the print head nozzle (±200-300µm).
To date, only one other study [70] has already encapsulated VIC aggregates (2*104 cells/
aggregate, formed by hanging drop) in HA-MA and HA-MA/Gel-MOD gels but only investigated
40
proliferation and migration of cells from the aggregates into the gel. In our study, viability,
proliferation, migration, fusion, phenotype, and ECM production of VIC in the encapsulated
micro-aggregates (10 and 20 w/v% Irgacure- and VA-crosslinked gels) were analyzed.
We have shown that VIC micro-aggregates in all conditions were viable up to 14 days.
However, VA-crosslinked gels disintegrated at earlier time points (7 and 14 days, respectively
for 10 and 20 w/v% gels) compared to Irgacure-crosslinked gels, which were stable for at least
14 days. A study by Hockaday et al. also reported the same problems: PEG-DA/Gel-MOD gels
with VA concentration under 0.5 w/v% dissolved when immersed in medium [73]. This could
suggest that the VA concentration used in our study was too low (20mol% VA=0.13 and 0.26
w/v%, for 10 and 20w/v% Gel-MOD respectively). Moreover, the gel fraction analysis showed
that VA-crosslinked gels (85.5-87.3%) had lower gel fraction percentage compared to the
Irgacure-crosslinked gels (90.5-97.5%). In other words, Gel-MOD material is less incorporated
in the 3D network of VA-crosslinked gels compared to the Irgacure-crosslinked gels.
We also noticed that the proliferation of VIC in the encapsulated micro-aggregates was less
pronounced compared to the non-encapsulated micro-aggregates and the number of Ki67+
cells decreased with increasing stiffness of the gels. The stiffness of the gels is correlated with
the concentration of Gel-MOD and the type of PI, as shown in our results that compressive
moduli of Irgacure-crosslinked gels ranged from 31.9 to 106.6 kPa and VA-crosslinked gels
ranged from 5.0 to 9.6 kPa. Previous studies have also described that the increase of Gel-
MOD concentration in the HA-MA/Gel-MOD hybrid gels led to a higher compressive modulus
of the gels [69, 70] and photopolymerization of PEG-DA/Gel-MOD hydrogels in the presence
of 0.025-0.1 w/v% Irgacure delivered higher mechanical stability than in the presence of 0.5-1
w/v% VA [73].
Moreover, it is noteworthy that the migration of VIC from the encapsulated micro-aggregates
into the gels was most pronounced in the 10 w/v% VA-crosslinked gels. Although not
investigated in this study, this migration could be due to the pore size of the gels. Benton et al.
have described that the pore formation in the gels is favorable for cell migration [60]. In
addition, previous studies have reported that the concentration of Gel-MOD and PI influence
the size, distribution, and interconnectivity of the pores [60, 76]. By scanning electron
microscopy it was visualized that heterogeneous, distributed, and interconnected pores were
found in the lower Gel-MOD concentration [76] and the largest pores were noticed in Gel-MOD
with the lowest Irgacure concentration [60]. Based on the proliferation and migration capacity
of VIC, which were more apparent in the soft gels, we also tested the fusion capacity of VIC
micro-aggregates in 10 w/v% VA-crosslinked gels by encapsulating a huge number of micro-
aggregates. We found that the encapsulated micro-aggregates fused within 3 days and bigger
41
valvular tissues were formed in various sizes (±150-1400µm). These valvular tissues also
showed time-dependent increase of GAG production and early reticular fibers production.
In general, we have shown that 10 and 20 w/v% Irgacure- as well as VA-crosslinked Gel-MOD
are suitable VIC micro-aggregate carrier material. However, the Gel-MOD properties can
influence the proliferation and the migration of VIC from the encapsulated micro-aggregates,
which was more pronounced in the softer hydrogels compared to the stiffer gels.
VIC in the encapsulated micro-aggregates should also be able to reflect the native VIC
phenotype and to synthesize ECM components and other bioactive molecules (MMP/TIMPs).
It has been previously shown that encapsulated VIC single cells in HA-MA/Gel-MOD hydrogels
were able to produce Col I, GAG, and elastin [70, 71, 74]. In this study, the ECM production of
VIC in the non-encapsulated and the encapsulated micro-aggregates was analyzed by
histochemical staining of gels from all conditions and RT-qPCR was performed at an early
and late time point to quantify the differences in VIC mRNA expression levels. For RT-qPCR,
micro-aggregates were encapsulated in 20 w/v% Irgacure-crosslinked Gel-MOD as we have
found that these gels remained stable for at least 21 days. Roosens et al. have already showed
that micro-aggregate culture can stimulate VIC mRNA expression of a broad panel of ECM
components [79]. This result was confirmed in this study, yet we also investigated the influence
of the encapsulation procedure as we compared mRNA expression levels of VIC in the non-
encapsulated micro-aggregates with VIC in the encapsulated micro-aggregates.
Collagen is the primary component of ECM in the heart valve leaflets and contributes to the
mechanical strength of the leaflets [80]. A native human valve leaflet consists of 74% Col I,
24% Col III and 2% Col V [9]. The mRNA expression levels of collagens (I, III, V) in VIC in the
encapsulated micro-aggregates increased over time, while remained stable in VIC in the non-
encapsulated micro-aggregates To visualize Col III, silver staining was used to detect the
presence of reticular fibers which are mainly composed of Col III in combination with other
collagen types (e.g. Col V), glycoproteins, and PG/GAG. Due to the silver impregnation,
reticular fibers stain dark [81]. The results showed an early and abundant reticular fiber
production of VIC in the non-encapsulated as well as in the encapsulated micro-aggregates
and the reticular fibers organized in circular bundles and aligned over time towards the edge
of the micro-aggregates. Col I immuno-staining was also performed to detect Col I production
in the encapsulated micro-aggregates. This staining showed that Col I fibrils were present and
increased over time (Fig. S3). Unfortunately, due to a production stop of the antibody, we
needed to switch to a new Col I antibody. This antibody has been optimized on a tissue sample
(porcine trachea), however, we believe that this new antibody reacts with other epitopes (it
also reacted with the gelatin hydrogel) and immature Col I fibrils in the encapsulated micro-
aggregates could not be shown. The middle layer of heart valve leaflet is mainly comprised of
42
GAG [7, 9]. Our study showed that the mRNA expression levels of HA, biglycan, decorin, and
versican in VIC in the encapsulated micro-aggregates increased over time and the expression
of HA, biglycan, and decorin was higher compared to the expression in VIC in the non-
encapsulated micro-aggregates. AB staining was also used to visualize GAG production. We
noticed that VIC in the encapsulated micro-aggregates showed more intense staining
compared to VIC in the non-encapsulated micro-aggregates. Another study has quantified
GAG content by a colorimetric assay and demonstrated that GAG content significantly
increased in Gel-MOD/HA-MA hybrid gels compared to pure HA-MA gels [70]. These findings
may suggest that Gel-MOD is likely to have stimulating capacity of GAG synthesis. Elastin is
predominantly distributed in the lamina ventricularis [1]. VIC in the encapsulated micro-
aggregates showed a time-dependent increase of elastin expression, while the expression of
elastin in VIC in the non-encapsulated micro-aggregates remained stable.
MMPs and TIMPs are associated with tissue/ECM remodeling. The synthesis and degradation
of ECM must be in balance to avoid abnormal tissue remodeling which has been observed in
heart valve diseases, including AS and AI [82]. In this study, VIC in the encapsulated micro-
aggregates showed a time-dependent increase in MMP-1/2/3/9, while TIMP-1/2 expression
remained stable over time. However, this increase was remarkably lower compared to the non-
encapsulated micro-aggregates.
In general, this study showed that Gel-MOD encapsulated VIC micro-aggregates, just like non-
encapsulated aggregates, were able to produce heart valve related ECM components and had
remodeling capacity.
One of the major challenges of working with VIC is the control of their activated/quiescent state
as VIC are mechanosensitive cells and are known to have an increased myofibroblastic
(activated) activity when cultured on stiff substrates, such as cell culture plastics [83-85]. The
transitional modulus between quiescent and activated state in monolayer cultured VIC is
approximately ≥15kPa [83]. Interestingly, our research group has shown that by cultivating VIC
in aggregates, VIC can regain their native, fibroblast-like quiescent phenotype [79]. This was
confirmed in our study but in addition, we also demonstrated that VIC in the encapsulated
aggregates remained in a quiescent state, as shown that α-SMA expression was low and FSP-
1 expression was high (20 w/v% Irgacure-crosslinked gels).
In contrast to the 2D cultured VIC that are activated on stiff substrates, the encapsulated VIC
single cells response differently to their substrate and show a reversed correlation between
the stiffness of the gels and VIC activation. Duan et al. have shown that α-SMA was more
expressed by VIC single cells in HA-MA gels (1.86 kPa) compared to VIC in HA-MA/Gel-MOD
(5.19 kPa) gels [70]. Mabry et al. have also demonstrated that dynamic stiffening of PEG-
based gels (from 0.24 to 13 kPa) reduced the activation level of the encapsulated VIC single
43
cells [86]. In addition, in the study of Hjortnaes et al. the encapsulated VIC single cells in HA/MA
Gel-MOD with a compressive modulus of 120 kPa showed low expression of α-SMA as well
[71]. These evidences obviously described that softer gels stimulated VIC to a more activated
phenotype as indicated by the increase of α-SMA and it has been reported that the fibroblast-
myofibroblast transitional modulus in 3D cultured encapsulated VIC single cells is 4.80–9.60
kPa [83] or ~7 kPa [84]. This phenotype transition is unexpected since softer gels more reflect
the in vivo mechanical properties of the native leaflets (±0.25-5 kPa [70]).
Duan et al. have shown that VIC single cells in HA-MA/Gel-MOD hybrid gels have multipotency
and are likely to undergo adipogenic- and limited chondrogenic- and osteogenic differentiation
when stimulated with differentiation media [72]. In this study, Sox-9, Runx2, Egr-1, and OCN
were analyzed since it has been showed that these markers can be elevated in calcific heart
valves [18-20]. Although OCN, Egr-1, and Runx2 (osteogenesis-related markers), were
expressed by VIC in the encapsulated micro-aggregates, there was no indication of
calcification. We also demonstrated that VIC in the encapsulated micro-aggregates expressed
significantly higher Sox-9 (early chondrogenic marker) compared to VIC in the non-
encapsulated micro-aggregates and to monolayer cultured VIC. This elevated expression of
Sox-9, and the time-dependent increase of GAG mRNA expression and AB staining could
indicate early spontaneous chondrogenesis of VIC.
44
CONCLUSION
In this study it is shown that VIC single cells and VIC micro-aggregates reacted differently
when encapsulated in gelatin hydrogels. Once encapsulated, VIC single cells showed a rapid
decrease in cell viability, while the encapsulated micro-aggregates did remain viable during
the complete study duration.
It is also shown that the concentration of Gel-MOD and the type of PI did not influence micro-
aggregate viability, but led to different gels with different stiffness. This gel stiffness influenced
proliferation and migration of VIC in the encapsulated micro-aggregates. 10 w/v% gels and
VA-crosslinked gels were softer compared to 20 w/v% gels and Irgacure-crosslinked gels. In
softer gels, VIC proliferation and migration from the micro-aggregates into the gels were more
pronounced, and VIC micro-aggregates were also capable to fuse and to form larger valvular
microtissues.
To ultimately produce a functional living tissue, micro-aggregates should mimic the native cell
phenotype and native tissue composition. It has been demonstrated that VIC in the
encapsulated micro-aggregates in all gels produced valve-related ECM components (GAG,
reticular fibers) and there was no indication of calcification. In addition, VIC from micro-
aggregates in Irgacure-crosslinked hydrogels showed downregulation of α-SMA- and
upregulation of FSP-1 expression, indicating stimulation towards a homeostatic quiescent
state. However, a significant increase of Sox-9 together with the massive GAG production
showed by AB staining and RT-qPCR could suggest an early chondrogenic differentiation of
VIC in these encapsulated micro-aggregates.
45
FUTURE PERSPECTIVES
There are some points that must be considered and further investigated when using Gel-MOD
encapsulated aggregates as a bio-ink for the generation of larger scale valvular tissues.
As shown in this study, the softer 10 w/v% VA-crosslinked gels allowed VIC from the
encapsulated micro-aggregates to proliferate, migrate, and fuse. However, the gel stability was
not optimal and needs to be improved, without affecting the current gel properties.
VIC phenotype in these softer hydrogels should also be investigated because so far we have
only analyzed VIC phenotype in micro-aggregates in the more stable and stiffer Irgacure-
crosslinked gels. Moreover, VIC phenotype markers should not only be analyzed on the RNA
level, but also on the protein level by performing (immuno)histochemical staining and western
blot.
Since the stiffer Irgacure-crosslinked gels showed a time-dependent increase of GAG
production, extremely elevated mRNA expression levels of PG/GAG, and increased
expression of the early chondrogenic marker SOX-9, it could be interesting to also investigate
late chondrogenic markers (aggrecan, Col II) in future soft-gel encapsulation experiments, to
rule out the possibility of chondrogenic differentiation in Gel-MOD hydrogels. Also, to confirm
the absence of calcium deposits, a Von Kossa staining can be performed.
In case the stability of VA-crosslinked gels can be improved and VIC in the encapsulated micro-
aggregates can remain in a homeostatic quiescent, undifferentiated state, VA-crosslinked gels
can be used to generate macro-scale valvular tissue. Otherwise, alternative hydrogels, such
as PEG-DA, HA-MA, alginate or hybrid gels, can be examined as a carrier material for VIC
micro-aggregates.
46
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i
SUPPLEMENT 1
S1: Morphology (BF photos and HE staining), viability (Ca/Pi staining), proliferation
(Ki67 immuno-staining), and ECM production (Col I, GAG, and Reticular fibers) of VIC
micro-aggregates from 1*106, 1.5*106, and 2*106 cells/microchip at day 3, 7, 14, and 21
Figure S1A: Morphology (BF microscopy and HE staining) of VIC micro-aggregates from 1*106, 1.5*106, and 2*106 cells/chip at day 3, 7, 14, and 21 [Scale bar = 200µm, magnification 10x (BF)], [Scale bar = 40µm, magnification 40x (HE)]
Figure S1B: Viability (Ca/Pi staining) and proliferation (Ki67 immuno-staining) of VIC micro-aggregates from 1*106, 1.5*106, and 2*106 cells/chip at day 3, 7, 14, and 21 [Scale bar = 200µm, magnification 10x (Ca/Pi)], [Scale bar = 40µm, magnification 40x (Ki67)]
ii
Figure S1C: ECM production (Col I, GAG, and reticular fibers) of VIC micro-aggregates from 1*106, 1.5*106, and 2*106 cells/chip at day 3, 7, 14, and 21 [Scale bar = 40µm, magnification 40x]
iii
SUPPLEMENT 2
Overview of studies that encapsulated VIC in hydrogels [References]
Author
(year)
Hydrogel/Crosslinking
Cell
type/Density/Study
duration
Results
[62]
Duan et al.
(2013)
Hybrid hydrogel: Gel-Alg
(0.06g/ml-0.05g/ml)
pVIC
hSMC
2*106 cells/ml
7d
Viability
Viability cells d1=91%; d7=86%; SMC d1=91%; d7= 85% Phenotype
Circularity: VIC d1=90%; d7=76%; SMC d1=91%; d7=80%
VIC expressed significantly higher vim than α-SMA (Fb-like), while SMC expressed significantly higher α-SMA than vim (MFb-like) Tensile Biomechanics
Ultimate TS and EM at d1 was significantly > in cell free hydrogels (TS=0.8MPa; EM=1.4MPa); at d7 values were statistically = between both groups (TS±0.35 MPa; EM±1 MPa)
Hydrogel discs (Ø8mm*1mm)
Crosslinking: 5m, 300mM CaCl2
[60]
Benton et al.
(2009)
10 w/v% Gel-MOD
DS =57%
pVIC
1*107 cells/ml
±5ng/ml TGF-β1
7w
Viability
Cells viable up to 7w (±80%); TGF-β1 did not influence viability Phenotype
Spreading and circularity (±90% and 62.5%) of TGF-β1 treated VIC significantly > compared to untreated (±50% and ±50%)
TGF-β1 lead to higher α-SMA and Col I gene expression Compressive Biomechanics
CM = 42kPa
Gels with lowest [PI] = largest pores (~1800µm2); gels with highest [PI]I = smallest pores (~400µm2)
Hydrogel discs (Ø5mm*1mm)
Crosslinking: 10m, UV-A (365nm), 5mW/cm2, 0.05-0.5 w/v% Irg
[69]
Duan et al.
(2014)
Hybrid hydrogels: HA-MA/Gel-MOD (2,4,6/6,10,12 w/v%)
DS = 22.5% HA-MA; 61.1% Gel-MOD
pVIC
5*106 cells/ml
7d
Viability
Cells viable up to 7d (>90%) Phenotype
Circularity VIC in gels with lower stiffness < gels with higher stiffness
d3 VIC in stiffer gels > proliferation than VIC in soft gels; d7 no significant difference in cell proliferation between gels; GAG content on d7>d3
VIC in stiffer gels expressed less α-SMA, vim, periostin; vim gene expression increased with decreasing gel stiffness; no difference in Col I expression of cells in different gels; upregulation Col I, vim, periostin and α-SMA after 7d
Compressive Biomechanics
CM decreased with increasing Gel-MOD concentration; CM increased with increasing HA-MA concentration
Highest stiffness: 4% HA-MA/6% Gel-MOD (13.0kPa); lowest stiffness: 6% HA-MA/12% Gel-MOD (4.2kPa)
Hydrogel discs (Ø8mm*1mm)
Crosslinking: 5m, UV-A (365nm), 2mW/cm2, 0.05 w/v% Irg
[70]
Duan et al.
(2013)
0.75 w/v% O0.1HA-MA or O0.05HA-MA or HA-MA
Hybrid hydrogels: + 1 w/v% Gel-MOD
DS = 22.5% HA-MA; 29% O0.05HA-MA; 30.7% O0.1HA-MA
pVIC
5*106 cells/ml
Spheroids = 5*106 cells/ml (hanging drop, 20µl)
14d
7d spheroid
Viability
Cells viable up to 14d (>75%); addition of Gel-MOD slightly increased cell viability Phenotype
At d14, cells more proliferative in softer hydrogels than in stiffer hydrogels; proliferation increased in hybrid hydrogels; circularity decreased in hybrid hydrogels and with decreasing CM
After 14d, GAG content in hybrid gels > than pure hydrogels; only significant increase of GAG content in HA-MA/Gel-MOD; α-SMA, periostin, and hyaluronidase I expression increased with decreasing stiffness of pure hydrogels; addition of Gel-MOD did not influence α-SMA expression, but vim expression slightly increased
Periostin and hyaluronidase expressions increased with decreasing stiffness of hybrid hydrogels
Cells from spheroids in hybrid hydrogels showed extensive spreading and fastest migration rate
HA-MA: α-SMA strongly expressed in VIC inside spheroids, migrated VIC showed more vim expression; Hybrid gels: cells inside spheroid more vim and less α-SMA, but migrated VIC showed more α-SMA
Compressive Biomechanics and Q
CM correlated with MW HA (HA-MA (2.74kPa) > O0.05HA-MA > O0.1HA-MA (1.86kPa))
Addition of Gel-MOD increased CM: O0.1HA-MA /Gel-MOD (5.19kPa) > HA-MA /Gel-MOD (3.93kPa)
Lower CM = higher Q; addition of Gel-MOD lowers Q (levels of crosslinks increased)
Hydrogel discs (Ø8mm*2mm)
Crosslinking: 5m, UV-A (365nm), 2mW/cm2, 0.03 w/v% Irg
iv
[71]
Hjortnaes et al.
(2015)
1 and 2 w/v% HA-MA
5 and 10 w/v% Gel-MOD
Hybrid hydrogels: HA-MA/Gel-MOD (1/5 w/v% and 2/10 w/v%)
DS = ±20% HA-MA; 80% Gel-MOD
pVIC
1*107 cells/ml
±5ng/ml TGF-β1
21d
Viability
Cells viable up to 21d (>80%); viability cells 2/10 w/v% gel (77.7%) < 1/5 w/v% gel (87.4%)
Amount of cells decreased in all gels from d1 till d21 except in 5w/v% Gel-MOD Phenotype
Cells in 5w/v % Gel-MOD spread the most; increased α-SMA, Col I, and MMP-9 expression (MFb-like); same levels of vim expression in all gels; Col I expression increased in all gels with Gel-MOD
aVIC in 5w/v % Gel-MOD; qVIC in hybrid- and HA-MA gels
TGF-β1 upregulated α-SMA, Col I, and MMP-9 expression in hybrid gels (qVIC/Fb-like to aVIC/MFb-like) Compressive Biomechanics and Q
1 compared to 2% HA-MA and 5 to 10% Gel-MOD showed 30% decrease in Q
CM increased with concentration of gel (5-120kPa); highest CM in hybrid hydrogels; CM inversely proportional to Q and pore size in independent hydrogels
Hydrogel discs (Ø8mm*450µm)
Crosslinking: 30-90s, UV-A (320-480nm), 2.5mW/cm2, 0.5 w/v% Irg
[72]
Duan et al.
(2015)
Hybrid hydrogels: HA-MA/Gel-MOD (0.5/1 w/v%)
DS = 22.5% HA-MA; 61.1% Gel-MOD
pVIC
hBM-MSC
hADSC
2.5*106 cells/ml
28d
Viability
All cell types viable up to 28d (>85%); Viability ADSC significantly > than BM-MSC Phenotype
Spreading and circularity VIC and ADSC > BM-MSC; at d28 VIC show the greatest spreading (lowest circularity)
GAG content of all cell types increased (d7-14), stable d14-d28; ADSC lowest GAG production
No significant changes in Col I content of VIC and ADSC; Col I content of BM-MSC peaked at d14; ADSC highest Col I
Elastin content of ADSC increased (d7-14) and stable d14-28; VIC and BM-MSC elastin production stable, but increased d28; ADSC highest elastin after d28
Hydrogel discs (Ø8mm*1mm)
Crosslinking: 5m, UV-A (365nm), 2mW/cm2, 0.05 w/v% Irg
[21]
Hjortnaes et al.
(2016)
Hybrid hydrogels: HA-MA/Gel-MOD (1/5 w/v%)
DS = ±20% HA-MA; 80% Gel-MOD
pVIC
21d
Viability
Viability cells d1=95%; d21=90% Phenotype
VIC quiescent in 3D hydrogels: low α-SMA Hydrogel discs (Ø8mm*450µm)
Crosslinking: 30s, UV-A (365nm), 2.5mW/cm2, 0.1 w/v% Irg
[86]
Mabry et al.
(2015)
MMP-degradable 8-arm PEGnb + adhesive peptide (RGDS)
pVIC
1*107 cells/ml
14d
Viability
Viability cells only evaluated after 2d: in soft gels ±90%; in stiff gels ±60% Phenotype
After 2d, only cell spreading in soft gels
Inverse relationship between gel stiffness and α-SMA/F-actin fibers expression; cells in stiffer gels more FSP-1 expression; VIC in soft gels MFb-like; VIC in stiff gel Fb-like; more MMP1 expression in stiff gels than soft gels; Col I expression is independent of gel-stiffness
Biomechanics
EM gels: 0.24 kPa - 13kPa
Hydrogel discs? (6mm mold)
Crosslinking: 3m, UV-A (365nm), 2mW/cm2, 1.7 mM LAP
[87]
Wu et al.
(2015)
4w/v% MMP-degradable PEG + adhesive peptide (RGDS)
pVIC
1*107 cells/ml
28d
±50µg/ml AA
Viability
Viability cells only evaluated after 3d=98% Phenotype
AA increased cell spreading by ±18%, MMP-2 production by 40%, and cell proliferation (amount of cells) by ±76% at d28 (compared to ±33% -AA)
Caspase 3/7 expression lower in AA-treated VIC than non-treated VIC; Col I and Col III production increased due to AA treatment; AA had no influence on α-SMA and ALP expression
Compressive Biomechanics
At d1 CM of cell-laden hydrogel < than cell-free hydrogel; at d21 or d28 CM of cell-laden hydrogel increased and = CM of cell-free hydrogel
CM: 4.3-6.0 kPa
Hydrogel discs (380µm-thick)
Crosslinking: 25s, white light exposure, 10μM eosin Y,1.5% v/v TEA and 3.5μl/ml NVP
[88]
Durst et al.
(2011)
5/10/15 w/v% of different MW PEG-DA pVIC
22*106 cells/ml
Phenotype
α-SMA expression decreased with increasing stiffness of gels
Hydrogel discs (6x44x0.88mm = WxHxT)
v
Crosslinking: 10m, UV-A (365nm), 10mW/cm2, 3 w/v% 2-hydroxy-40-(2-hydroxyethoxy)-2-methylpropiophenone
21d
Compressive and Flexure Biomechanics
Cell-free hydrogels stiffer than cell-laden hydrogels
[89]
Gould et al.
(2016)
2.5 mM MMP-degradable 8-arm PEGnb + adhesive peptides (RGDS/VGVAPG/P15)
DS > 95%
pVIC
15*106 cells/ml
42d
Viability
Viability cells after 14d = 95% regardless the adhesive peptides Phenotype
VIC metabolic activity in RGDS-gel slightly > than other gels at d14
At d14, VIC in RGDS-gel more elongated than in VGVAPG/P15-gels
At d2, VIC in RGDS-gels had slightly higher MMP activity compared to VGVAPG/P15-gels; cell spreading appeared to correlate with MMP activity: cell spreading by d14 in P15 gels + increased MMP expression
At d14 VIC α-SMA+ RGDS-gel (33%), VGVAPG-gel (56%) and P15-gel (38%); VIC in VGVAPG-gels < Col I but > Col X and CS than RGDS-gel
After 42d, Col deposition in RGDS-gels > than VGVAPG-gels and P15-gels; Col X/ Col I ratio in RGDS-gels < than VGVAPG-gels and P15-gels
Compressive Biomechanics
CM = 2kPa
Crosslinking: 10m, UV-A (365nm), 8.5mW/cm2, 0.2 mM Irg
[90]
Shah et al.
(2008)
Hybrid hydrogels: HA-MA/PEG-diMA (1.38µM/1.37 to 22.0mM)
DS ~ 85%
pVIC
22*106 cells/ml
6w
±1.2mg/ml Exogenous HA
±1U/ml HAase
Viability
HA enhanced VIC survival up to 2w at concentration > than 1.38µM Phenotype
Col production in HAase-untreated gels only around the cells, while in HAase-treated gels better Col distribution away from the cell body, throughout the gel and cell number increase
Elastin production was apparent after 4w in HAase-treated gels; no apparent elastin staining in the untreated gels
Exogenous HA reduced Col production in HAase-treated gels compared to untreated gels and increased elastin production in both gels Compressive Biomechanics
Total macromere mass increased with increasing PEG concentration; dry mass HA decreased with increasing PEG concentration
Increasing PEG concentration caused decrease in Q (Q = 12.4-37.3)
Hydrogels discs (Ø5mm*1mm)
Crosslinking: 7m, UV-A (365nm), 5mW/cm2, 0.0375 w/v% Irg
[59]
Masters et al.
(2005)
2 w/v% HA-MA
10 w/v% PEG-DA
Hybrid hydrogels: HA-MA/PEG-DA (2/5 w/v% and 2/10 w/v%)
DS = 1.5%, 54%, 100% HA-MA
pVIC
40*106 cells/ml
6w
±5 HA-MA degradation products
Viability
VIC were viable after one week of encapsulation Phenotype
Addition of HA degradation products influenced VIC proliferation; VIC proliferation increased with decreasing MW HA degradation products
At d3, total matrix production of VIC (+ degraded HA) was similar to the control gels; at d20 there was significant difference in total matrix production and composition
At d3, elastin production of VIC (+ degraded low MW HA) increased compared to untreated control cells; at d20, increase of elastin more apparent in almost all gels supplemented with HA of different MW
Encapsulated VIC showed significant elastin production after 6w Compressive Biomechanics
Q decreased with increasing MA; CM and degradation time decreased with increasing degree of crosslinking
Highest Q: 2% HA-MA (79.9); lowest Q: 2% HA-MA + 10% PEG-DA (9.69); Highest CM: 2% HA-MA (12.61 kPa); lowest CM: 2% HA-MA + 10% PEG-DA (54.7)
Crosslinking: 3m, UV-A (365nm), 5mW/cm2, 0.05 w/v% Irg
[91]
Benton et al.
(2009)
5 and 10 w/v% MMP-degradable 4-arm PEGnb + adhesive peptide (RGDS)
pVIC
15*106 cells/ml
14d
±5ng/ml TGF-β1
Viability
VIC still viable after 14d in both gel concentrations; 5 w/v% gels more viable cells than 10 w/v% gels Phenotype
VIC were proliferative and showed random migration in gels
VIC in 5 w/v% gels spread faster (less circular) than VIC in 10 w/v% (14d)
VIC spread more in gels with higher RGD concentration
+ TGF-β1= increased α-SMA and Col I expression
Hydrogels discs (Ø7mm*1mm)
Crosslinking: 10m, UV-A (352nm), 5mW/cm2, 0.05 w/v% Irg
[73]
Hockaday et al.
(2017)
Hybrid hydrogels: 20 w/v% Gel-MOD/ PEG-DA (ratio = 1:1) + 7.5 w/v% alginate
hVIC, hADSC, hSMC
Viability
Highest viability for SMC= 95%, VIC= 93%, ADSC= 93%
Irg had a more detrimental effect on the viability of all cell types than VA-086 Hydrogels discs (Ø8mm*1.5mm)
vi
Crosslinking: 5min, UV-A (365nm), 2mW/cm2, 0.025-0.1 w/v% Irg, 0.25-1 w/v% VA-086
(also high powered LED lamp = 136mW/cm²)
2.5*106 cells/ml
7d
Catalase
Increased PI concentration = decreased viability Phenotype
ADSC spread more than VIC and SMC in Irg gels; VIC circularity increased with increasing Irg; Cells in VA-086 gels spread more
Catalase did reduce oxidative stress but did not improve viability of cells in general
Using LED lamp resulted in decreased numbers of viable cells Compressive Biomechanics
CM of hydrogels ranged from ±19.3-120.7 kPa; CM of VA-086 and Irg crosslinked gels with the low intensity light was not significantly affected by increasing PI concentration; lamp light intensity seemed to have saturation points since increased PI concentrations did not increase the CM
Gels with 0.25-0.5 w/v% VA-086 mostly dissolved; crosslinked gels with 0.75-1 w/v% VA-086; 0.025-0.1 w/v% Irg = ok
CM of gels with 1.0 w/v% VA-086 (26kPa) and 0.1 w/v% Irg (80kPa) increased when using LED to ±50 and ±125 kPa respectively
Q decreased when increasing PI concentration and increasing light intensity (only for VA-086)
[74]
Duan et al.
(2016)
Hybrid hydrogels: HA-MA/Gel-MOD (0.5/1 w/v%)
DS = 22.5% (HA-MA); 61.1% (Gel-MOD)
hVIC
4*106 cells/ml
21d
± activation medium (TGF-β)
± osteogenic medium
Viability
Viability of the cells did not significantly decrease during the study Phenotype
In activation media 3kPa VIC-laden gels (5m crosslinking) had higher GAG, Col, and elastin content than 1kPa VIC-laden gels (1m crosslinking) at d7. At d21 in osteogenic media 1kPa gels had higher GAG, Col, and elastin content than 3kPa gels; Gels in control media = Col and elastin in 3kPa gels higher than in 1kPa gels,
Expression of α-SMA and vim detected in all hydrogels after 21d; VIC in activation or osteogenic media more pronounced α-SMA+
Increasing gel stiffness resulted in upregulation of α-SMA and downregulation of vim in activation media as well as in osteogenic media (significant in 3kPa gels)
Runx2 and OCN detected only in VIC in osteogenic media and increased with increasing stiffness; VIC in 1kPa hydrogels showed more significant increase of Runx2 and OCN than in 3kPa gels; ALP expression increased with increasing stiffness (higher in 1kPa gels)
Compressive Biomechanics
CM correlated with crosslinking duration; the stiffness of acellular hybrid hydrogels increased from ±1.33kPa (1m) to ± 3.33kPa (5m)
Stiffness of VIC-laden gels (5m crosslinking) decreased at d1; after 21d VIC-laden gels in control media maintained stiffness, while VIC-laden gels in activation and osteogenic medium increased stiffness; gels in osteogenic media had the highest stiffness
Stiffness of VIC-laden gels (1m crosslinking) similar to stiffness of acellular gels; only VIC-laden gels in osteogenic media increased stiffness; stiffness of gels in osteogenic media from 1m crosslinking had much greater increase in stiffness than gels from 5m crosslinking
Hydrogels discs (Ø 8mm*1mm and 15mm*15mm)
Crosslinking: 1 or 5m, UV-A (365nm), 2mW/cm2, 0.05 w/v% Irg
[76]
Chen et al.
(2013)
Microfluidic: 5, 10, 15 w/v% Gel-MOD
DS = 94%
Co-culture: pVIC-pVEC
1.5*106 cells/ml
21d
Viability
Cell viability after 4d = 84%
Concentration of Gel-MOD had no significant influence on the cell viability; cell viability decreased with increasing PI concentration (94% viability with 0.5% PI; 0% viability with 2%PI)
Phenotype
VIC in 5 w/v% gels had the highest spreading; cell spreading decreased with increasing PI concentration and with increasing UV irradiation time
After 3d, VEC formed a confluent monolayer; VEC remained viable and did not migrate
Static cultured VIC = 78% α-SMA+, static co-cultured VIC = 29% α-SMA+, shear-stressed co-cultured VIC 17% = α-SMA+) Compressive Biomechanics
EM ranged from 2.5kPa (5 w/v% gels) to ~30kPa (15 w/v%); pore size decreased with increasing gel concentration (20µm to 7µm); pores in 5 w/v% gels are heterogeneous in size and distribution, and interconnected; pores in 15 w/v% gels were uniform in size and distribution, and separated by thin walls
Channel (2000*300µm = WxH)
Crosslinking: 30s (5 w/v% gel), 25 s (10 w/v% gel), and 15s (15 w/v% gel), UV-A (360-480nm), 7.2mW/cm2, 0.5, 1, 2 w/v% Irg
[92]
Zhang et al.
(2015)
5, 10, 20 w/v% MMP-degradable PEG-DA + adhesive peptide (RGDS)
DS >95%
pVIC
3*107 cells/ml
28d
Viability
Cell viability >80% in all conditions; cell viability decreased with lower amount of peptide Phenotype
Increasing the amount of RGDS resulted in decrease of cell circularity and increase cell area
VIC activation = higher level of α-SMA found in gels with higher RGDS; amount of RGDS had little effect on vim
PEG hydrogels were degraded by VIC; MMP2 was detected in culture medium
Hydrogels discs (Ø 8mm*1mm)
Crosslinking: 3m, UV-A (360-480nm), 7.2mW/cm2, 0.3 w/v% Irg
vii
Crosslinking: 3m, white light exposure, 1.5 w/v% TEA, 10 µM Eosin Y, 0.375 w/v% NVP
Compressive Biomechanics
CM dependent on crosslinking time; Q of 10 w/v% dependent on PEG-DA MW (lower Q for stiffer 3.4kDa gels, higher Q for weaker 20kDa gels
EM of cell-laden hydrogels was significantly higher at d21 and 28 than at d1 to d14
[93]
Mabry et al.
(2016)
8-arm MMP-degradable PEGnb + adhesive peptide (RGDS, IKVAV, RLD)
DS = ~90%
pVIC
1*107 cells/ml
5d
Viability
After 1d cell viability = 75% Phenotype
VIC in control gels (without peptides) remained round after 5d; cells in IKVAV gels remained round, while cells in RGDS or RLD gels were elongated; circularity of cells decreased with increasing peptide concentration
VIC in gels without RGDS had low levels of α-SMA; VIC in gels with RGDS were more activated after 3d (MFb-like) Compressive Biomechanics
EM= 1 or 6 kPa
Hydrogels discs (Ø4mm*1mm)
Crosslinking: 2m, UV-A (365nm), 2.5mW/cm2, 1.7mM LAP
Abbreviations: p = porcine; h = human; d = days; w = weeks; m = minutes; s = seconds; VIC = valvular interstitial cells; SMC = smooth muscle cells; α-SMA = alpha smooth muscle; vim = vimentin; Fb = fibroblasts; MFb
= myofibroblasts; MOD = modified; Gel = gelatin; MA = methacrylamide; Alg = alginate; TS = tensile strength; EM = elastic modulus; DS = degree of substitution; CM = compressive modulus; HA = hyaluronic acid; Abs =
absorbance; GAG = glycosaminoglycans; Col = collagen; OHA = oxidized hyaluronic acid; Q = swelling ratio; Irg = Irgacure 2959; MW = molecular weight; PI = photo-initiator; BM-MSC = bone marrow derived mesenchymal
stem cells; ADSC = adipose derived mesenchymal stem cells; MMP = matrix metalloproteinase; PEG = polyethylene glycol; Hep = heparin; PEGDA: polyethylene glycol diacrylate; PEGnb = polyethylene glycol norbornene;
TEA = triethanolamine; NVP = N-vinylpyrrolidone; LAP = lithium phenyl-2,4,6-trimethylbenzoylphosphinate; W = width; H = height; T = thickness; CS = chondroitin sulphate; PEG-diMA = polyethylene glycol dimethacrylate;
HAase = hyaluronidase; CS = chondroitin sulfate
viii
SUPPLEMENT 3
S3: Col I immuno-staining with the old antibody on VIC micro-aggregates in 10 and 20
w/v% Irgacure-crosslinked Gel-MOD at day 1, 3, and 7
Figure S3: Col I production detected by the old Col I antibody in 10 and 20 w/v% Irgacure-crosslinked Gel-MOD at day 1, 3, and 7 [Scale bar = 40µm, magnification 40x]
10% Gel-MOD
20% Gel-MOD
D1 D3 D7