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The interplay between biomaterial degradation and tissueproperties : relevance for in situ cardiovascular tissueengineeringCitation for published version (APA):Brugmans, M. C. P. (2015). The interplay between biomaterial degradation and tissue properties : relevance forin situ cardiovascular tissue engineering. Technische Universiteit Eindhoven.

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The interplay between biomaterial

degradation and tissue properties

Relevance for in situ cardiovascular tissue engineering

Marieke Brugmans

A catalogue record is available from the Eindhoven University of Technology Library ISBN: 978-90-386-3848-5

Copyright © 2015 by M.C.P. Brugmans All rights reserved. No part of this book may be reproduced, stored in a database or retrieval system, or published, in any form or in any way, electronically, mechanically, by print, photo print, microfilm or any other means without prior written permission by the author. Printed by Ipskamp Drukkers B.V., Enschede, the Netherlands. The research and printing of this thesis was supported by:

Financial support by the Dutch Heart Foundation for the publication of this thesis is gratefully acknowledged. This work was supported by a grant from the Dutch government to the Netherlands Institute for Regenerative Medicine (NIRM, grant No. FES0908).

The interplay between biomaterial degradation and

tissue properties


PROEFSCHRIFT

ter verkrijging van de graad van doctor aan de Technische Universiteit Eindhoven, op gezag van de

rector magnificus, prof.dr.ir. F.P.T. Baaijens, voor een commissie aangewezen door het College voor

Promoties in het openbaar te verdedigen op woensdag 10 juni 2015 om 16.00 uur

door

Maria Cornelia Philomena Brugmans

geboren te Veghel

Dit proefschrift is goedgekeurd door de promotoren en de samenstelling van de promotiecommissie is als volgt: Voorzitter: prof. dr. P.A.J. Hilbers

1e promotor: prof.dr.ir. F.P.T. Baaijens

2e promotor: prof.dr. C.V.C. Bouten

copromotor: dr. A. Driessen-Mol

leden: dr. J. Kluin (UvA)

dr. P. Habibovic (UM)

dr.rer.nat. C. Ottmann

adviseur: dr. P.Y.W. Dankers

I

Contents

Summary III

Chapter 1: General introduction 1

1.1. Human cardiovascular tissues

1.1.1 Heart valves

1.1.2 The heart valve leaflets

1.1.3 Blood vessels

1.2 Cardiovascular diseases and current treatments

1.3 Cardiovascular tissue engineering approaches and challenges

1.4 Biomaterials

1.4.1 Natural biomaterials

1.4.2 Synthetic biomaterials

1.5 In vivo resorption of biomaterials

1.5.1 Resorption pathways

1.5.2 Variation in resorption of biomaterials

1.6 The host response to biomaterials

1.6.1 The phases of the natural healing response

1.6.2 Macrophage phenotypes

1.7 Rationale and outline

2

2

3

4

5

7

10

11

11

12

12

13

14

14

15

16

Chapter 2: Polycaprolactone scaffold and reduced in vitro cell culture:

Beneficial effect on compaction and improved valvular tissue formation

19

2.1 Abstract

2.2 Introduction

2.3 Materials and Methods

2.4 Results

2.5 Discussion

2.6 Conclusion

20

21

23

26

33

38

Chapter 3: Superior tissue evolution in slow-degrading scaffolds for valvular

tissue engineering

39

3.1 Abstract

3.2 Introduction


3.4 Results

3.5 Discussion

3.6 Conclusion

40

41

42

44

50

53

II

Chapter 4: Hydrolytic and oxidative degradation of electrospun supramolecular

biomaterials: In vitro degradation pathways

55

4.1 Abstract

4.2 Introduction


4.4 Results

4.5 Discussion

4.6 Conclusion

56

57

59

62

68

72

Chapter 5: Advanced electrospun scaffold degradation by inflammatory

macrophages in comparison with healing macrophages

73

5.1 Abstract

5.2 Introduction


5.4 Results

5.5 Discussion

5.6 Conclusion

74

75

76

81

87

88

Chapter 6: General discussion 91

6.1 Main findings of the thesis

6.2 Towards the most promising tissue engineering approach and scaffold material

6.3 Study limitations and the future of in-situ cardiovascular tissue engineering

6.4 Conclusion

92

96

101

104

References 107

Nederlandse samenvatting 127

Dankwoord 129

Curriculum vitae 131

List of publications 133

III

Summary

The interplay between biomaterial degradation and tissue properties:


Various tissue engineering (TE) approaches are currently under investigation to create

cardiovascular tissue replacements. The most promising strategy is the in-situ TE

approach, in which off-the-shelf available synthetic electrospun scaffolds are used to

replace diseased vessels or heart valves. After implantation, a host inflammatory response

is activated, leading to the infiltration of macrophages, which play a key role in both

scaffold degradation and tissue formation. As a result, a living tissue that is able to

remodel and adapt to the environmental changes is obtained in-situ. It is crucial to select

the optimal scaffold material to ensure mechanical integrity immediately after

implantation, which starts degrading as soon as sufficient tissue is formed to take over

the native function. The aim of the research described in this thesis was to examine the

interplay between scaffold degradation rates and the amount and composition of the

formed tissue within the scaffold. Furthermore, degradation characteristics of scaffolds

manufactured from different supramolecular biomaterials, were investigated.

By imbalance between scaffold degradation and tissue formation, the mechanical

integrity cannot be ensured and compaction and retraction of in-vitro TE heart valves

occurs, causing regurgitation in-vivo. We studied whether compaction could be reduced

by the use of slow-degrading polycaprolactone (PCL) instead of fast-degrading poly-4-

hydroxybutyrate coated polyglycolic acid (PGA-P4HB) electrospun scaffolds and/or the

use of a lower cell passage number to enhance tissue formation. The use of slow-

degrading materials improved resistance to retraction of TE valvular leaflets and reduced

compaction of TE rectangular scaffold strips. In addition, tissue formation, stiffness, and

strength increased with decreasing cell passage number, but did not affect compaction of

the engineered tissues.

Thereafter, the effect of scaffold degradation rate on the amount and composition of

tissue, the mechanical integrity, and the tissue to scaffold ratio were investigated. Slow-

and fast-degrading scaffolds were seeded with vascular cells or kept unseeded. We

hypothesized that the cells in fast-degrading scaffolds would compensate for the rapid

loss of mechanical integrity by increased tissue production. Increasing amounts of tissue

with time were shown in both scaffold groups, which was indeed more pronounced for

PGA-P4HB-based tissues during the first two weeks of culture. Ultimately, PCL-based

tissues resulted in the highest amount of tissue after 6 weeks. In addition, we described a

method to correct for the amount of remaining scaffold weight, in order to allow a fair

comparison between in-vitro engineered tissues grown on scaffolds with a different

Summary

IV

degradation rate and in-vitro engineered tissues and native tissues. By implementation of

this correction, extracellular matrix values similar to values of native pulmonary heart

valves were found. The amounts of collagen crosslinks were still below native values in all

engineered tissues, but did display a continuing increase during culture.

In-vivo, degradation of scaffold materials can be accomplished by the (enzymatic

accelerated) hydrolytic and/or the oxidative pathway. To investigate both pathways,

separately and in an accelerated fashion, in-vitro degradation assays were designed. For

in-situ TE of cardiovascular tissues, the supramolecular materials PCL-2-ureido-[1H]-

pyrimidin-4-one (PCL-UPy) and PCL-bisurea (PCL-BU) are used, due to their combination

of strength and elastic properties. Degradation characteristics and susceptibility to the

hydrolytic or the oxidative degradation pathway of these materials were investigated and

compared with those of conventional PCL. Depending on the morphological and chemical

composition of the materials, conventional and supramolecular PCL-based scaffolds

responded differently to both degradation pathways. Conventional PCL is more prone to

hydrolytic enzymatic degradation as compared to the supramolecular materials, while the

opposite was shown when degraded by the oxidative pathway. We demonstrated the

ability of tuning degradation characteristics by mix-and-match PCL backbones with

supramolecular moieties. This allows screening and selecting the optimal biomaterial for

pre-clinical studies targeted to different clinical applications.

Macrophages are known to play an important role in the degradation of the implant,

however the contribution of macrophage phenotype to scaffold degradation was still

unclear. The inflammatory phenotype is known to secrete both reactive oxygen species

(ROS) and enzymes involved in scaffold degradation. However, degradation might also be

accomplished by the healing phenotype, as also these secrete enzymes involved in

scaffold degradation. The correlation between macrophage phenotype and degradation

of electrospun scaffolds was investigated in this thesis. We elucidated that the

macrophage phenotype affected the contribution to scaffold degradation, consolidating

that inflammatory macrophages indeed accelerated degradation. In addition, the

electrospun PCL induced macrophage polarization towards the healing phenotype, which

is a beneficial feature for in-situ TE.

In conclusion, the choice of scaffold material is of high importance to maintain mechanical

integrity. Results in this thesis emphasize that a slow-degrading material is favored over a

fast-degrading material, as mechanical integrity will be maintained for a longer period,

which is important for in-situ TE purposes. Furthermore, tissue seemed better organized

when cultured on slow-degrading scaffold materials and therefore is less prone to

compaction. In addition, this thesis demonstrated that degradation characteristics can be

tailored, which is essential as different degradation characteristics are desired for various

applications.

1

C

General introduction

1

Chapter 1

2

Cardiovascular in situ tissue engineering using synthetic materials is a promising approach

for overcoming the limitations of current available treatments for the repair or

replacement of damaged or diseased cardiovascular tissues. Off-the-shelf, highly porous

scaffolds, in the shape of the desired construct, act as templates for replacing the diseased

tissues with healthy tissues. They should provide temporary mechanical strength, while

endogenous cells are attracted to the implanted scaffold and produce new tissue. At the

same time new tissue is formed, the implanted synthetic scaffold slowly resorbs and is

ultimately removed from the body, leaving behind functional, viable tissue that is able to

adapt to environmental changes. In order to maintain good mechanical integrity,

immediately after implantation until the newly formed tissue takes over this role, the

balance between tissue formation and bioresorption of the scaffold is of high importance.

As a consequence, bioresorption of the implanted scaffold plays a crucial role in the final

outcome of the engineered construct. Tunability of scaffold resorption in vivo is desired,

as different resorption rates are needed for various applications. Furthermore, amount

and quality of newly formed tissue might be influenced when growing in slow- or fast-

resorbing materials. In vitro, no resorption by the body occurs and therefore the break

down of scaffolds in this thesis is referred to as degradation. The aim of this thesis is to

elucidate in vitro degradation characteristics of scaffolds manufactured from different

synthetic biomaterials, and the effect of degradation on tissue formation. With the use of

this knowledge, bioresorbable scaffolds can be created with appropriate resorption

characteristics for use as cardiovascular tissue replacements.

1.1 Human cardiovascular tissues

1.1.1 Heart valves

The heart is a muscular organ that regulates blood flow throughout the body in order to

transport oxygen and nutrients to tissues, and remove metabolic waste from the tissues.

Oxygen-poor blood returns back into the right ventricle, via the right atrium. When the

right ventricle contracts, this blood is pumped through the pulmonary artery into the

lungs, where it becomes oxygenated again. The left ventricle receives this oxygen-rich

blood from the lungs via the left atrium. Upon contraction of the left ventricle, blood is

pumped into the aorta and distributed throughout the whole body. To ensure

unidirectional blood flow, the heart is provided with four valves: the tricuspid valve, the

mitral valve, the pulmonary valve, and the aortic valve (Figure 1.1). The tricuspid and

mitral valves (atrioventricular valves) are situated between the atria and the ventricles,

and prevent blood from flowing back from the ventricles into the atria. The pulmonary

and aortic valves (semilunar valves) are situated between the right ventricle and the

pulmonary artery, and the left ventricle and the aorta, respectively. These valves prevent

blood from flowing back from the pulmonary artery and aorta into the ventricles. Valves

open and close approximately 100.000 times each day and about 3.7 billion times in a


3

1 lifetime, subjecting the thin and flexible leaflets to loads and deformations with every

heartbeat.

Figure 1.1 Schematic transverse sections of the human heart and its four valves. Cross section of the heart, anterior view (A). Direction of blood flow is indicated with arrows. Cross section of the heart, top view, showing the opened (B) and closed (C) position of the pulmonary and aortic valves to allow blood flow from the ventricles into the pulmonary artery and aorta (adapted from zoominmedicine.com).

1.1.2 The heart valve leaflets

The pulmonary and aortic valves are referred to as semilunar valves due to the half-moon

shape of their three thin leaflets. The leaflets are connected to a fibrous, ring shaped

thickening of the arterial wall, called the annulus. Leaflets are composed of cells,

embedded in an extracellular matrix (ECM). The cross section in Figure 1.2 shows that

leaflets have a layered architecture, which comprise three distinct layers; the fibrosa, the

spongiosa and the ventricularis. These layers can be identified in both the aortic and

pulmonary heart valve leaflets, however, a more pronounced fibrosa layer can be found

in the leaflets of the aortic heart valve.

Figure 1.2 Cross section of one of the leaflets of a heart valve (left). Schematic overview of the composition of a leaflet, consisting of three distinct layers, each comprising valvular interstitial cells (VICs) and ECM components (right) (adapted from Vessely, 1998 and Schoen, 2013).

Chapter 1

4

Each layer has a specific composition and organization of the ECM. The fibrosa, which is

located at the arterial side of the leaflet, consists mainly of a dense collagen network and

provides mechanical strength to the tissue. The spongiosa, situated between the fibrosa

and the ventricularis, consists mainly of proteoglycans and water-binding

glycosaminoglycans (GAGs) to absorb shocks on the leaflet. The ventricularis, at the

ventricular side, is rich in elastin fibers, which ensure flexibility and restores the

contracted configuration of the leaflets [1, 2]. Two types of cells are present within the

leaflets; valvular interstitial cells (VICs) and valvular endothelial cells (VECs). VECs form a

single endothelial layer, covering the whole leaflet surface to prevent direct contact of the

ECM with blood, and thereby provide a non-thrombogenic layer. VICs are the most

abundant cellular component of the heart valves and are found throughout the leaflets.

In healthy adult heart valves, these cells reside in a fibroblast-like quiescent state, but they

can differentiate into myofibroblasts-like cells and mediate ECM synthesis and remodeling

[3-7].

1.1.3 Blood vessels

There are three major types of blood vessels; the arteries, the veins and the capillaries

(Figure 1.3). In general, arteries carry oxygen- and nutrient-rich blood away from the

heart, after which the actual exchange of oxygen and nutrients between blood and tissues

takes place in the capillaries. Oxygen-poor blood is collected in the veins, and is carried

back to the heart. Capillaries consist of only a single layer of endothelial cells to enable

optimal gas and nutrient exchange.

Figure 1.3 Schematic overview of the human circulatory system. The further away vessels are from the heart, the smaller they become. In the smallest vessels, the capillaries, nutrient and oxygen exchange takes place (From Martini, Frederich H.; Timmons Michael J.; Tallitsch, Robert B.; Human anatomy, 7th Edition, © 2012. Reprinted by permission of Pearson Education, New York).


5

1 Within the arteries and veins three different layers of tissue can be distinguished; the

tunica adventitia, the tunica media and the tunica intima. These layers differ in thickness,

depending on the function of the vessel. The tunica adventitia, which is the outer layer,

consists of loosely woven collagen fibers and may contain nutrient capillaries in the larger

vessels. In the middle layer, the tunica media, smooth muscle cells and elastin can be

found, which regulate and assist in vasodilatation or vasoconstriction. The tunica intima,

situated at the lumen, is in direct contact with blood and consists of a single endothelial

layer and some elastic fibers [8].

1.2 Cardiovascular diseases and current treatments

Cardiovascular diseases (CVD) remain the leading cause of death worldwide, among both

men and women, resulting in almost half of all deaths in Europe and one third of all deaths

in the United States [9, 10]. Among CVD, coronary artery disease is the most frequent and

is often treated with bypass grafting [9, 10]. Each year, surgeons perform approximately

800.000 coronary bypass surgeries worldwide [11]. Furthermore, vascular grafts are

needed in diabetic patients, end-stage renal disease and pediatric heart operations.

Autologous small-diameter arteries and veins are the preferred replacement grafts [12]

despite 50% of the grafts occluding within 10 years [13]. However, it is estimated that

these tissues are not available in 30% of all patients, due to either inherent disease or

harvest during previous operations [14]. In these cases, non-degradable synthetic grafts

can be used (Figure 1.5), such as expanded polytetrafluoroethylene (ePTFE, i.e. GORE-

TEX®) or polyethylene terephthalate (PET, i.e. Dacron®). These materials are widely used

in the clinic for over 50 years and have shown to be successful for medium to large

diameter vascular graft applications. However, data on small-diameter grafts (<6mm) is

still very poor. Results showed that these synthetic grafts are prone to thrombus

formation and intimal hyperplasia, leading to occlusion of the graft [12, 15, 16]. Therefore,

they have lower patency rates compared to autologous grafts, with patency rates of 24-

44% for PTFE compared to 70% for saphenous veins after 5 years in peripheral

applications [17]. This shows the obvious need of small-diameter vascular grafts that

resemble autologous grafts.

Heart valve disease (HVD) can occur in any single valve, or a combination of several valves.

Diseases related to the aortic and mitral valves are most common and result in the highest

mortality rate, because of its important hemodynamic positions [9, 18]. HVD can lead to

stenosis (narrowing of the valve opening), or regurgitation (leakage of the valve) (Figure

1.4). These pathologies can be caused by a congenital abnormality (e.g. 2 leaflets instead

of 3), calcification or by damage to the valve due to rheumatic fever or endocarditis [18-

21].

Chapter 1

6

Figure 1.4 Schematic cross sections of a healthy heart during systole (A) and diastole (C). During systole, the left ventricle contracts and opens the aortic valve, allowing blood flowing into the aorta. In case of a stenotic valve (B), blood flow is obstructed, resulting in thickening of the left ventricle. During diastole, the left ventricle relaxes and fills with blood. The aortic valve is closed, to prevent backflow from the aorta. Regurgitation, due to incomplete closed leaflets (D), results in an enlarged heart cavity and a thickened left ventricle (adapted from Nishimura 2002 with permission from Wolters Kluwer Health).

The most common treatment of end-stage disease is replacement of the valve.

Worldwide, approximately 290.000 heart valve replacements are performed each year,

and this number is expected to increase up to 850.000 by 2050, due to aging of the

population and the increased ability to diagnose valvular heart disease [9, 22]. Current

available heart valve replacements are either mechanical or bioprosthetic (Figure 1.5),

each having their own benefits and disadvantages [23-26]. Different mechanical valves

have been designed; ball-and-cage valves, mono-leaflet valves and bi-leaflet valves, which

are made of for example carbon, Teflon or titanium [25]. Mechanical valves can last a life-

time, with a valve replacement rate <2% over 25 years [27], and are readily available.

However, they are prone to thrombus formation due to non-physiological flow profiles

that result in blood cell damage [28]. As a consequence, life-long anti-coagulation therapy

is required, which results in increased risk of spontaneous bleeding in those patients.

Bioprosthetic valves can be harvested from a human (homograft) or from an animal

(xenograft). Homografts are closest to natural valves and can be derived from a donor

(allograft) or from patients themselves (autograft). Donor valves are sterilized using anti-

biotic and anti-fungi solutions and stored by cryopreservation or fixated. However, there

is limited availability of this type of valves. Xenografts, made of glutaraldehyde fixed

porcine or bovine material are often used instead. These valves do not require anti-

coagulation therapy, but are prone to tissue degeneration and calcification, with

reoperation rates of 20% after 10 years and 30% at 15 years, which limits their durability

[25, 29-31]. Furthermore, the risk of transmission of animal diseases to human and

immunogenic reactions is increased with this type of valve replacement [32, 33]. An

alternative is to decellularize these tissues, and thereby decrease the immunological

response without limiting the remodeling capacity of the implants [15, 34]. This results in

native-like tissue replacements, which can be implanted as such [35], or can be re-seeded

with autologous cells prior to implantation [36-38].


7

1

Figure 1.5 Examples of existing heart valve and vascular replacements. Starr-Edwards prostheses (ball-cage) (A), Medtronic open pivotTM mechanical valve (B), Edwards SAPIEN 3 Transcatheter Heart Valve (C), Medtronic Hancock II® bioprostheses (D), Medtronic Melody® transcatheter pulmonary valve (E), Medtronic Contegra® pulmonary valved conduit (F), GORE-TEX® vascular grafts (G), Dacron® vascular grafts (H). Images A and C are reproduced with permission of Edwards Lifesciences LLC, Irvine, CA. Images B, D, E, and F are reproduced with permission of Medtronic, Inc., a subsidiary of Medtronic plc. Image G courtesy of W. L. Gore & Associates, Inc.

Although the current available cardiovascular tissue replacements significantly improve

quality of life and life expectancy, a shortcoming is that they are not able to adapt to

changing physiological demands, as they consist of non-living materials. The development

of a living tissue that can adapt is of utmost importance to further improve quality of life

and life expectancy of patients with cardiovascular diseases.

Growth potential has also been assumed as a desired property of a living heart valve

replacement to prevent re-operations in pediatric patients. As current treatments do not

accommodate for growth, oversized replacements are often used in pediatric patients to

prevent early outgrowth of the replacement. However, research on failed replacements

in children have shown that not outgrowth of the replacement, but contracture and

stenotic valves are the most important failure modes [39-41]. This indicates that

preventing the most common failure modes should have priority over growth potential.

1.3 Cardiovascular tissue engineering approaches and challenges

Different cardiovascular tissue engineering approaches are used within the field of

regenerative medicine. These include the classical in vitro tissue engineering approach,

with or without decellularization of the created tissues afterwards, the in vivo tissue

Chapter 1

8

engineering approach, and the promising in situ tissue engineering approach (Figure 1.6).

They all aim to create a living cardiovascular substitute that is able to adapt after

implantation.

The classical approach is the in vitro tissue engineering approach, in which autologous

cells, to prevent immune responses, are expanded in vitro. After cell seeding, the

bioresorbable scaffold construct is often subjected to stimuli, which mimic physiological

pressures and/or flows in a bioreactor to enhance tissue formation [42-45]. Different cell

sources have been examined [4] including vascular-derived cells [46], which are also used

in our lab [47-49], neonatal cell sources [50-52], mesenchymal stem cells [53, 54], adipose-

derived cells [55], and endothelial progenitor cells [56, 57]. Also, different materials are

used to create scaffolds, which include the natural polymers e.g. fibrin and collagen [15],

and the synthetic polymers e.g. PCL [58]. Weinberg and Bell produced the first tissue

engineered vascular graft (TEVG) based on collagen and vascular-derived cells, using this

in vitro approach [59]. However, it was found that this graft was mechanically unstable

and not suitable for implantation. Encouraging progression was made within this field, as

shown in both in vitro and in vivo studies on vascular grafts, with high patency rates up to

13 months [60-65]. However, to date no living small-diameter vascular graft is made that

remains patent during a life-time. The in vitro tissue engineering approach has also shown

to be promising for clinical applications. Engineered tubes based on a bioresorbable

scaffold material, seeded with autologous vascular cells, have been successfully implanted

into humans to reconstruct the pulmonary artery [46, 66]. Proof of concept of an in vitro

tissue engineered heart valve (TEHV) was demonstrated in 1995 by Shinoka et al. [67],

where a single autologous tissue engineered leaflet was implanted in a sheep. The next

step was to develop functional three-leaflet tissue engineered valves. This was first

reported by Hoerstrup, Sodian and Stock. They showed functionality of TEHV at the

pulmonary position for up to 24 weeks [68-70]. More recent studies also showed

promising in vivo results of TEHV with functional leaflets in sheep for up to eight months

[53]. Nevertheless, the main challenge in all recently performed pre-clinical studies is

retraction of the heart valve leaflets leading to regurgitation [43, 71-73]. The balance

between the contractile tissue-producing cells and the mechanical integrity of the

remaining scaffold is very important in order to prevent this retraction of the leaflets [47,

74]. Therefore, researchers decellularize the tissue-engineered constructs after in vitro

culture in order to remove the contraction forces exerted by these cells. Furthermore, this

decellularization protocol is used to create off-the-shelf available tissue replacements [43,

73, 75-77]. While decellularization of native tissues has demonstrated various rates of

repopulation after implantation, decellularization of in vitro cultured constructs has

shown faster host cell repopulation [73]. This is probably due to the lack of elastin barriers

and a less mature collagen structure [78-80].

In vivo tissue engineering can be defined as a process where the peritoneal cavity or

subcutaneous space is used to generate an autologous graft by taking advance of the


9

1 immune response to foreign materials. This in vivo formed tissue can subsequently be

removed and used as a vascular graft [62, 81-85]. The main challenge remains to maintain

all vascular grafts patent after implantation.

Figure 1.6 Overview of the different tissue engineering approaches. In the in vitro tissue engineering technique, cell-seeded scaffolds are placed into bioreactors, to mature the tissue before implantation (A.1, middle and right photos made by Bart van Overbeeke) or are decellularized before implantation (A.2, reprinted from Dijkman 2012 with permission from Elsevier). Scaffold in the shape of a blood vessel or a heart valve is implanted directly in the in situ tissue engineering approach (B). In vivo tissue engineering makes use of an e.g. silicon rod which is implanted into the peritoneal cavity. After some time this rod is explanted and the tissue formed around this rod is used for replacement of the diseased tissue (C, reprinted from Yamanami 2013 with permission from Springer).

Although tissues with properties towards autologous grafts can be created with the in

vitro and in vivo tissue engineering approaches, it takes weeks to months to produce these

implants. Together with regulatory issues for transportation and storage of tissue

engineered implants, this approach is very expensive and time-consuming. Furthermore,

it could result in products with batch-to-batch variation in tissue quality due to variation

in performance of biological material.

To circumvent these disadvantages, a trend towards in situ tissue engineering is seen in

academic research and industry [86, 87]. Within this approach, a synthetic bioresorbable

scaffold is either implanted cell-free [88-92], or pre-seeded with autologous cells prior to

implantation [72, 93-96]. The scaffold, in the shape of the desired replacement, should

maintain mechanical functionality immediately after implantation, while endogenous

cells are attracted to the implanted material and produce new tissue. While neo-tissue is

Chapter 1

10

formed, the scaffold slowly degrades and is ultimately removed from the body, leaving

behind living tissue that is able to grow and adapt to changing physiological demands.

From nearly 30 years ago until now, encouraging results of in vivo studies using cell-free

vascular grafts in rats, dogs and the canine model have been reported [88, 90, 91, 97-101].

Further, promising data of clinical trials based on in situ tissue engineering of large

diameter TEVG, is reported. Large diameter bioresorbable vascular grafts, pre-seeded

with autologous bone marrow mononuclear cells before implantation into pediatric

patients, demonstrated growth capacity, while no graft-related mortality or graft failures

were observed during a mean follow-up of 5.8 years [46, 94, 95, 102]. The unguided in

situ tissue engineering process in pristine scaffolds, where no cells, proteins, or other

biologicals are added to the scaffold before implantation, is here referred to as

endogenous tissue growth (ETG). Recently, the company Xeltis implanted cell-free,

bioresorbable vascular conduits into five pediatric patients. These conduits were designed

to enable ETG and resulted, to this date, in successful tissue replacements [92]. In situ

tissue engineering of heart valves also showed good progress during the last years.

Pulmonary valves, based on a bioresorbable material and pre-seeded with autologous

bone marrow cells, were implanted into non-human primates, and demonstrated a

confluent layer of endothelial cells after 4 weeks and proper valvular functionality up to 4

weeks [72]. In a recent study performed by the Dutch BioMedical Materials program

‘iValve’, cell-free heart valve constructs, based on a bioresorbable scaffold, were

implanted at the pulmonary position in an ovine model. After 12 months, functional heart

valve leaflets were demonstrated with good tissue formation [103].

In conclusion, several tissue engineering approaches have demonstrated promising

results, although each approach still has challenges to overcome. The in situ tissue

engineering approach is especially appealing and promising. This ‘device-based’ approach

is based on faster, easier, and cheaper production of off-the-shelf available grafts and

encounters less regulatory hurdles, compared to cell-based approaches. Of particular

interest is the ETG approach, where a bare scaffold is used without any additives.

1.4 Biomaterials

Selection of the right biomaterial is important within tissue engineering, as mechanical

integrity should maintain until neo-formed tissue can take over this role. This is mainly

important for in situ tissue engineering, as a bare scaffold is implanted, which should

provide sufficient mechanical strength by itself, immediately after implantation. The ideal

biomaterial for cardiovascular tissue engineering should also be highly porous with an

interconnected pore network to allow for cell infiltration and tissue in-growth, nutrient

supply, and removal of metabolic waste products. Furthermore, it should be

biocompatible, bioresorbable, reproducible, and contain mechanical properties that are

consistent with the anatomical site of implantation to prevent compliance mismatch.


11

1 Different types of biomaterials, either natural or synthetic, are described within this

section.

1.4.1. Natural biomaterials

Apart from native matrices that are decellularized before implantation, scaffolds could

also be made from natural materials. These include fibrin, elastin, hyaluronan, silk fibroin

and collagen [15, 77, 104, 105], which show good biocompatibility in terms of chronic

inflammation and toxicity, and closely mimic the natural ECM of tissues. A disadvantage

of these materials is the high batch-to-batch variations and researchers have limited

control, although progression is made, over the material properties, which often results

in lack of mechanical performance [15, 106, 107].

1.4.2. Synthetic biomaterials

Bioresorbable synthetic biomaterials are widely used for cardiovascular tissue

engineering purposes. They are cheap to fabricate, readily available and researchers have

better control over critical properties, such as the resorption rate or mechanical

properties compared to the natural biomaterials. Among them are PCL, polyglycolic acid

(PGA) and polylactic acid (PLA), which are used in medical devices that are already

approved by the Food and Drug Administration (FDA) or have European Conformity (CE)

mark registration [58, 108]. Recent studies showed promising results for cardiovascular

applications with these and other polymers, like polyglycerolsebacate (PGS) and

polyurethanes [73, 91, 109, 110]. Each polymer has different characteristics in terms of

mechanical properties or resorption and might be suitable for different applications. PCL

has been shown to be an interesting candidate for TEVG [111], however, due to its limited

fatigue resistance, this material is less suitable for TEHV, as in TEHV the materials are

exposed to demanding mechanical loads.

A unique and new set of synthetic materials are the supramolecular polymers, which are

formed by arrays of directed, non-covalent interactions, such as hydrogen bonds, between

the polymer chains (Figure 1.7) [112, 113]. These supramolecular polymers can form

complex 3D-structures by self-assembly. Material properties such as mechanical

properties and/or resorption rate can be modified easily by combining or changing ratios

of the same building blocks, providing a broad variety of biomaterial properties. As the

monomeric units in supramolecular materials possess relatively low molecular weights,

they can easily be dissolved and processed. Examples of the supramolecular biomaterials

are the PCL-based materials modified with 2-ureido-[1H]-pyrimidin-4-one (UPy) [114-118]

or bis-urea (BU) [119] units. These exhibit strong and elastic properties and therefore

might be more suitable for cardiovascular applications like heart valves, when compared

to some of the conventional polymers, e.g. PCL.

Chapter 1

12

Figure 1.7 Schematic overview of an example of a supramolecular PCL-based material with ureidopyrimidinone moieties (grey blocks). Polymer chains are held together via hydrogen bonds (dotted lines).

1.5 In vivo resorption of biomaterials

Within the field of tissue engineering, the balance between tissue formation and scaffold

resorption, which is different in every application, is of high importance. Bioresorbable

scaffolds should provide mechanical strength until sufficient mature neo-tissue is formed

to take over this function.

1.5.1. Resorption pathways

In vivo, implanted scaffolds can be degraded by different pathways that may operate at

the same time and that even may affect each other (Figure 1.8). These are the hydrolytic

and the oxidative resorption pathways. During hydrolysis, chemical bonds (mostly esters)

of the polymer chain are cleaved by the reaction of water molecules, forming shorter

polymer chains and finally oligomers or monomers that can be cleared from the body

[120, 121]. Enzymes, like esterases, which are present in the blood or are secreted by

macrophages and other activated cells after implantation of the scaffold, are known to

accelerate this process [122]. For example, lipases are known to accelerate PCL resorption

[123-125].

The oxidative resorption pathway is mediated by reactive oxygen species (ROS) that are

secreted by inflammatory cells, like macrophages, neutrophils and giants cells, that are

recruited to the scaffold fibers [124, 126]. These ROS include hydrogen peroxide (H2O2),

nitric oxide (NO), hydroxyl radical (·OH) and superoxide (O2-) and are responsible for chain

scission of the polymers [127, 128]. Previous studies have investigated that oxidation of

polymers is often initiated by abstraction of a hydrogen atom by radicals, resulting in chain

scission of the polymer [127].


13

1 Resorption can arise by two different mechanisms; surface erosion or bulk erosion [122].

In surface erosion, the exterior layer of the material is affected, while the core remains

intact until the surrounding layer has been resorbed. This typically results in mass loss,

thinning of the material and a stable molecular weight of the inner part of the material.

Bulk erosion occurs throughout the whole material simultaneously, resulting in decreased

molecular weight and mass loss throughout the material.

Figure 1.8 Schematic overview of in-vivo resorption pathways. After implantation, cells attach to the

scaffold fibers and secrete both enzymes and ROS. This results in resorption of the scaffold fibers via the

enzymatic accelerated hydrolytic pathway, and/or the oxidative pathway. Depending on the chemical

composition and the morphology of the biomaterial, one of these pathways plays a more dominant role.

Here, material A is affected by enzymatic hydrolysis, resulting in thinning of the fibers (surface erosion),

while fibers are unaffected by the oxidative pathway. Material B is unaffected by the enzymatic pathway,

but demonstrates broken fibers (bulk erosion) due to the ROS products generated in the oxidative pathway.

[Drawing courtesy from Anthal Smits]

1.5.2. Variation in resorption of biomaterials

Resorption properties of widely used materials such as polyesters, polyethers and

polyurethanes have been examined extensively. The mechanism and rate of material

resorption depend on environmental factors, such as temperature, pH and mechanical

stress [122]. Furthermore, the chemical composition and morphology of the polymers

have an influence on the resorption rate [106]. In general, it is shown that polymers

containing ester bonds react with water molecules and undergo hydrolysis. PGA is a

hydrophilic material, in which water molecules can enter easily, resulting in fast hydrolytic

resorption [129, 130]. The polyester PCL is a more hydrophobic material and results in

slower hydrolytic resorption compared to PGA [129]. Solutions are able to be in contact

with a larger surface area of the material, often resulting in faster resorption, when a

porous scaffold is created compared to a solid film. Other polymers, including

polyurethanes, were found to be more susceptible to the oxidative resorption pathway

[127, 131].

Chapter 1

14

1.6 The host response to biomaterials

Physiological wound healing is a response to injury, induced by the implantation of a

biomaterial. This healing response involves complex, well-regulated processes, which

include the four overlapping phases of haemostasis, inflammation, proliferation and

remodeling (Figure 1.9). The entire healing response is mediated by cytokines and growth

factors, which are secreted by different cell types involved in this host response.

1.6.1. The phases of the natural healing response

Phase 1: Hemostasis (seconds to minutes)

After the first interaction of the biomaterial with blood, proteins from the blood and

interstitial fluid adsorb to the biomaterial, dependent on the biomaterial surface

properties [126, 132, 133]. These proteins serve as binding sites for leukocytes [134].

Platelets also adhere to the biomaterial and secrete chemoattractants for immune cells

that are involved in the second inflammatory phase.

Phase 2: Acute inflammation (minutes to days)

During the early phase of acute inflammation, the most prominent cell type that migrates

from the blood toward the implanted biomaterial are neutrophils. After 24-48 hours,

these neutrophils undergo apoptosis and are phagocytosed by resident tissue

macrophages. Monocytes enter the site of implantation and differentiate into

macrophages. Macrophages function as phagocytic cells that clear wound debris and cell

remnants, or foreign material.

Phase 3: Proliferation (days to weeks)

After 3 to 5 days, fibroblasts, which are recruited by macrophages, enter the site of

implantation and start to deposit ECM proteins like fibronectin, collagen and

proteoglycans. Furthermore, new blood vessels are generated by endothelial cells within

the newly formed tissue during this regeneration phase.

Phase 4: Remodeling (weeks to years)

During the remodeling phase, there is clearance of macrophages. The final outcome of

tissue regeneration or scar formation is dependent on the duration of the chronic

inflammatory phase. In case of an optimal healing process, the scaffold is completely

degraded and phagocytosed by the cells, while ECM is synthesized, matured, and

remodeled simultaneously. In case of a prolonged healing response, fibrous scar tissue

will be formed. It is believed that foreign body giant cells play an important role in this

prolonged healing response, as they continuously activate fibroblasts, resulting in

excessive deposition of ECM components [132]. This often results in encapsulation of the


15

1 (remaining) scaffold by avascular fibrous connective tissue instead of complete resorption

of the scaffold and full replacement by native-like tissue [135].

Figure 1.9 The four phases of wound healing, which include hemostasis, inflammation, proliferation and remodeling. Different cell types are involved in each phase, which undergo apoptosis when they fulfilled their function. Figure adapted from Enoch and Leaper 2008.

1.6.2. Macrophage phenotypes

Upon migration into affected or inflamed tissue, monocytes differentiate into

macrophages. Dependent on micro-environmental signaling factors, macrophages can

polarize into a heterogeneous population with different markers and functions. Different

classes of macrophages, based on these markers and functions, have been identified and

are believed to play an important role in the balance and final outcome of tissue

regeneration or scar formation. The classically activated, pro-inflammatory macrophages

are referred to as the M1 type. These are activated by pro-inflammatory signals, such as

interferon gamma (IFN-ɣ) and lipopolysaccharide (LPS), and secrete pro-inflammatory

cytokines and ROS. M2 type macrophages are the alternatively activated, anti-

inflammatory cell type, involved in immunoregulation and wound-healing [132]. These

cells are activated by molecular cues such as IL-4 and IL-13. Others have described more

subsets of the M2 macrophages, called M2a, M2b and M2c. M2a and M2b are both

associated with wound healing and immunoregulatory functions, while M2c is involved in

suppression of the immune response [136-139]. Although the classes of macrophage

phenotypes are defined, it is well known that these classes are the extremes of a

continuum, and the macrophage phenotype is plastic and can change due to micro-

environmental factors [136]. In an optimal healing process, the macrophages should

undergo a phenotypic change from the M1 type during the inflammation phase towards

the M2 type during the regenerative phase. Furthermore, several studies suggest that the

Chapter 1

16

scaffold surface, fiber diameter and pore size might also influence the macrophage

phenotype [140-145]. This indicates that scaffold fiber morphology and composition

should be carefully selected to promote optimal healing responses.

1.7 Rationale and Outline

One of the main challenges of tissue engineering is to control the balance between tissue

formation and in vivo scaffold resorption. In order to design a scaffold with appropriate

resorption properties for cardiovascular in situ tissue engineering applications, the aim of

this thesis is to elucidate in vitro degradation characteristics of scaffolds manufactured

from several (supramolecular) biomaterials and the effect of degradation rates on tissue

formation and composition.

A disturbed balance between tissue formation and scaffold degradation, where scaffolds

degrade too fast in combination with traction forces exerted by the cells, resulted in

compaction and retraction of in vitro tissue engineered heart valves, causing regurgitation

in vivo when these valves were implanted [43, 71-73]. Therefore, we used in vitro tissue

engineering in chapter 2, to study whether the use of 1) slow- (PCL) instead of a fast-

degrading (PGA-P4HB) electrospun scaffold meshes and 2) a lower cell passage number

to enhance tissue formation, has beneficial results on compaction. Furthermore, tissues

were engineered using both ovine and human cells, to determine the effect of

interspecies differences on tissue development.

In our search for the appropriate scaffold material for cardiovascular applications, we

investigated and discussed in chapter 3 how tissue development and composition

changed during 6 weeks of in vitro culture, when cells were cultured on slow- (PCL) and

fast-degrading (PGA-P4HB) electrospun scaffolds. Furthermore, the values of ECM

components and collagen crosslinks were measured in the tissue engineered constructs

and compared to values found in native human heart valves.

To take another step forward in the world-wide availability of cardiovascular grafts, in situ

tissue engineering seems to be a very promising alternative, as this approach results in a

reduction in costs, production time, and regulatory issues related to tissue culture,

compared to the classical in vitro tissue engineering approach. This also induces other

demands, e.g. prolonged mechanical integrity, on the scaffold material, as the grafts are

implanted as bare scaffolds, without any pre-cultured tissue. Supramolecular materials

like PCL-UPy or PCL-BU are promising for in situ tissue engineering as these materials

comprise strong and elastic properties, which are desired properties to replace load-

bearing tissues like heart valves. To better understand the degradation characteristics of

various electrospun scaffolds, accelerated in vitro degradation assays were designed in

chapter 4. With the use of these assays, the degradation characteristics of different


17

1 electrospun supramolecular materials were explored and compared to a conventional

material.

Macrophages are known to play an essential role in the resorption of the implanted

scaffold meshes. To illustrate whether there is a correlation between the inflammatory

(M1) or healing (M2) macrophage phenotypes and degradation of electrospun meshes, in

vitro culture systems were used in chapter 5. In addition, we investigated the preferred

polarization phenotype of macrophages when cultured onto PCL meshes with a fiber

diameter of 10 µm.

Finally, the main findings of the thesis are summarized and discussed in chapter 6. This

includes a discussion on remaining challenges and the required future (research) steps

towards safe clinical application of cardiovascular tissue replacements.

Chapter 1

18

19

Poly-ε-caprolactone scaffold and

reduced in vitro cell culture:

Beneficial effect on compaction and

improved valvular tissue formation

2

M. Brugmans

A. Driessen-Mol

M. Rubbens

M. Cox

F. Baaijens

J Tissue Eng Regen Med

2013

Chapter 2

20

2.1 Abstract

Tissue-engineered heart valves (TEHV), based on PGA scaffold coated with poly-4-

hydroxybutyrate (P4HB), have shown promising in vivo results in terms of tissue

formation. However, a major drawback of these TEHV is compaction and retraction of the

leaflets causing regurgitation. To overcome this problem the aim of this study was to

investigate 1) the use of the slow degrading PCL scaffold for prolonged mechanical

integrity and 2) the use of lower passage cells for enhanced tissue formation.

Passage 3, 5 and 7 (p3, p5 and p7) human and ovine vascular-derived cells were seeded

onto both PGA-P4HB and PCL scaffold strips. After 4 weeks of culture, compaction, tissue

formation, mechanical properties and cell phenotypes were compared. TEHV were

cultured to observe retraction of the leaflets in the native like geometry.

After culture, tissues based on PGA-P4HB scaffold showed 50-60% compaction, while PCL-

based tissues showed compaction of 0-10%. Tissue formation, stiffness and strength were

increased with decreasing passage number, however, this did not influence compaction.

Ovine PCL based tissues did render less strong tissues compared to PGA-P4HB based

tissues. No differences in cell phenotype between the scaffold materials, species or cell

passage numbers were observed.

This study shows that PCL scaffolds may serve as alternative scaffold material for human

TEHV with minimal compaction and without compromising on tissue composition and

properties, while further optimization of ovine TEHV is needed. Reducing cell expansion

time will result in faster generation of TEHV, providing a more rapid treatment to patients.

PCL scaffold and reduced in vitro cell culture

21

2

2.2 Introduction

With an increasing number and aging of the world population, valvular heart disease is an

expanding health problem. Approximately 290 000 heart valve replacements are

performed annually worldwide and this number is estimated to increase up to 850 000 by

2050 [22]. Bioprosthetic and mechanical heart valves, which are successfully used for

decades, improve quality of life and life prolongation for most patients [146, 147].

However, these valves have some restrictions as they consist of non-living and non-

autologous materials. Therefore, they are not able to grow, adapt or remodel to changing

physiological environments, resulting in decreased durability [22]. Furthermore,

bioprosthetic valves are susceptible to calcification, while mechanical valves require

lifelong anticoagulation therapy to prevent thrombo-embolism [22, 146]. To overcome

these problems, researchers are studying the possibility of creating tissue engineered

heart valves (TEHV) [146]. Patients’ own cells are incorporated, resulting in valves of

autologous living tissue that are able to grow, remodel and adapt to the changing

environment after implantation[146]. Our approach to create such TEHV is to isolate

patients’ cells from the vena saphena magna, expand them in vitro up to the desired

amount of cells and subsequently seed them onto a bioresorbable synthetic scaffold in

the shape of a heart valve. After a culture period in a bioreactor of 4 weeks, where the

valves are exposed to mechanical stimuli in order to stimulate tissue formation, the valves

are able to withstand systemic pressures in in vitro tests [148], aiming ultimately at

implanting them into patients.

Different types of synthetic scaffolds are used for cardiovascular tissue engineering

applications. In particular, PGA scaffold, coated with P4HB, or combined with another

scaffold material, showed to be a promising candidate in terms of tissue formation, as

demonstrated in vascular graft studies [42, 60] and in vivo TEHV studies [68, 71, 72, 149].

Hoerstrup et al. demonstrated in an ovine model that after 20 weeks in vivo, the valves

yielded an organized, layered structure with many architectural features and ECM

characteristics that are present in native valves. In vivo, PGA and P4HB are resorbed

completely within 4 and 8 weeks respectively [68]. The downside of using this rapid

resorbing PGA scaffold is compaction (flattening of the leaflets) and retraction (shrinkage

of the leaflets), causing regurgitation [71, 72, 150]. This is a result of traction forces

exerted by the cells, likely in combination with an imbalance of the newly formed tissue

and loss of mechanical integrity of the scaffold due to degradation [74, 148, 151]. Rabkin-

Aikawa demonstrated TEHV containing αSMA positive cells during in vitro culture, while

after 20 weeks in vivo, there was a strong decrease of αSMA positive cells [6]. As αSMA is

related to traction forces of the cells [152], we assume that after 20 weeks, these traction

forces will be decreased in vivo. Therefore, a scaffold with proper mechanical integrity

during in vitro culture and the first months after implantation is desired to withstand the

cell traction forces during this phase. The use of a slower degrading scaffold material such

Chapter 2

22

as PCL may represent a promising alternative, as TEHV can be produced that are

mechanically reliable for months, thereby offering sufficient mechanical integrity to

prevent tissue compaction and retraction [153]. As PCL can be processed by

electrospinning, it is possible to create complex geometries and mold the scaffold directly

into the desired 3D shape of a heart valve [153]. This direct 3D molding is not feasible for

PGA scaffolds, which are only available in sheets. Another benefit of PCL is the possibility

to create thin leaflets with a thickness of 300 µm, while the PGA meshes are produced

with a thickness of 1000 µm. As PGA-P4HB scaffolds are more rapidly degrading, the cells

might be exposed to larger magnitudes of mechanical loading as compared to the cells in

PCL scaffolds, which might on their turn be partly protected from loads by the long-term

presence of the scaffold. As the stress level exerted on the vascular cells is known to

change phenotype of the cells towards activated myofibroblasts[154], tissue formation

capacity of cells in the two scaffold types might differ, along with different phenotypes

[155-158]. Therefore, it is important to compare cell phenotype, tissue formation capacity

and compaction in tissues based on both scaffold types when considering the use of PCL

as a scaffold material to produce TEHV. Based on the above, we hypothesize that the cells

in PGA-P4HB might have a more activated phenotype accompanied by increased tissue

formation capacity as compared to cells in PCL scaffolds.

Another alternative to tackle compaction and retraction of TEHV might be by using cells

of a low passage number. Aging cells, due to in vitro expansion, lose their potential to

proliferate [159, 160]. Currently in our lab, cells are expanded up to passage 6-7 to ensure

enough cells for seeding multiple TEHV [148]. Whether the amount of tissue formation or

cell phenotype in 3D cultures is influenced by the use of cells of a low passage number is

still unclear as to the best of our knowledge, previous work on the effect of cell aging by

expansion has been performed on 2D cultures only. Therefore, the role of cell aging in 3D

tissue formation capacity needs to be further investigated. We hypothesize that cells of a

low passage number (passage 3) are more productive, resulting in more tissue formation

and of a higher quality, compared to cells of a high passage number (passage 7). This

improved tissue formation capacity on its turn may result in less compaction and

retraction, as it is influencing the balance between matrix quality and the mechanical

integrity of the scaffold towards increased matrix quality. We assume that the increased

matrix formation will increase the resistance to the traction forces exerted by the cells.

An additional benefit of using cells of a lower passage number is the reduction in cell

expansion time, which will result in faster generation of TEHVs and, thereby, providing a

more rapid treatment to patients.

To summarize, the aim of this study is to evaluate alternative approaches to overcome

the compaction and retraction of TEHV as observed with the use of rapid degrading PGA-

P4HB scaffolds, without compromising on tissue composition and properties. The

alternative approaches that are being studied here are 1) the use of a slow degrading PCL


23

2

scaffold for prolonged mechanical integrity and 2) the use of lower passage vascular cells

for enhanced tissue formation. Compaction, tissue formation, cell phenotype and

mechanical properties of engineered tissues based on passage 3, 5 and 7 vascular cells in

both PCL and PGA-P4HB scaffolds are compared. TEHV aim to be designed for humans,

but since the ovine model is commonly used to show proof of principle, both human and

ovine cells were used.

2.3 Materials and methods

2.3.1 Cell culture

Human vascular-derived cells were harvested from segments of a vena saphena magna

from a 60 years old patient that underwent bypass surgery, and was obtained according

to the Dutch guidelines for secondary used materials. Ovine vascular-derived cells were

obtained from the vena jugularis of adult sheep of approximately 2 years old (n=2,

Swifter). The cells were isolated via the outgrowth method. In short, endothelial cells of

the vessels were removed by incubation with a collagenase solution. Remaining

endothelial cells were removed from the lumen side using a cell scraper. After removal of

the endothelial cells, the vessels were minced into small pieces of approximately 1 mm2

and the fragments were plated into 6 wells-plates. The outgrowing cells were expanded

using standard culture methods in a humidified atmosphere containing 5% CO2 at 37°C,

and passaged at 90-100% confluency.

Plating densities were 3.3-4.6*103 per cm2 for human and 1.6-2.3*104 per cm2 for ovine

cells, based on differences in cell size. Isolation and expansion medium consisted of

advanced Dulbecco’s Modified Eagle Medium (DMEM; Invitrogen, Breda, Netherlands),

supplemented with 1% GlutaMax (Invitrogen), 1% Penicillin/Streptomycin (P/S, Lonza,

Basel, Switzerland), and 10% Fetal Bovine Serum (FBS, Greiner Bio one, Frickenhausen,

Germany) for human cells or 10% Lamb Serum (Invitrogen) for ovine cells. During culture,

cells of all passage numbers grew in the characteristic ‘hill and valley’ morphology,

indicating smooth muscle cells.

2.3.2 Scaffold preparation and sterilization

Rectangular strips (25x5 mm) were cut out of PGA meshes (PGA, Cellon, Bascharage,

Luxemburg) and conventionally electrospun PCL meshes, with a thickness of 1000 μm and

300 μm, respectively. As heart valves contain a more complex geometry compared to

strips, which might result in differences in terms of compaction, trileaflet heart valve

scaffolds were fabricated using scaffold meshes of the same thickness. PGA scaffolds were

additionally coated with poly-4-hydroxybutyrate (P4HB, received via a collaboration with

Prof. Hoerstrup of the University Hospital Zurich) to provide structural integrity to the

Chapter 2

24

mesh. The outer 3-4 mm of both PGA and PCL scaffold strips were attached onto stainless

steel rings (RVS Paleis, Geleen, Netherlands) using 15% polyurethane-tetrahydrofuran

(PU, Desmopan) glue, leaving an 18*5 mm area for cell seeding. The solvent was allowed

to evaporate overnight in a vacuum oven. PCL scaffolds were sterilized by gamma

irradiation (Isotron, Ede, Netherlands). PGA-P4HB scaffold sterilization was achieved by

immersion in 70% sterile ethanol for 30 minutes. To facilitate cell attachment, the

scaffolds were incubated overnight with tissue engineered (TE) medium, consisting of

expansion medium supplemented with 0.25 mg/ml L-ascorbic acid 2-phosphate (Sigma).

Lamb serum (0.1%) and FBS (10%) was added to ovine and human TE medium,

respectively.

2.3.3 Cell seeding and tissue culture

Passage 3, 5 and 7 (referred to as p3, p5 and p7) cells were seeded onto both PGA and

PCL scaffolds (n=6 per passage and scaffold for each cell type), with a seeding density of

20 million cells per cm3 using fibrin as a cell carrier [161]. In short, cells were suspended

in TE medium containing thrombin (10 U/ml, Sigma). This cell suspension was mixed with

an equal volume of TE medium containing fibrinogen (10 mg/ml, Sigma) and dripped onto

one side of the scaffolds before polymerization of the gel was accomplished. As control

strips, three PGA and PCL scaffolds were seeded with fibrin only. After seeding, the

scaffolds were placed in an incubator at 37°C for 30 minutes, to allow polymerization of

the fibrin gel. Thereafter, 6 ml of TE medium was added to each scaffold.

The strips were cultured for 4 weeks and TE medium was changed twice a week. For the

heart valve cultures, passage 7 cells were used and seeded according to similar protocols

as for the strips. After seeding, the valves were placed in a bioreactor system and cultured

for 4 weeks [43] .

2.3.4 Compaction

Compaction was assessed from upper view photographs of the strips that were taken

once a week. The valves were photographed after 4 weeks only. Compaction of the strips

was defined as the reduction of width, compared to the width at the start of culture.

Photographs were analyzed using the program Image J (version 1.43u).

2.3.5 Biochemical assays

For the quantification of tissue formation after 4 weeks of culture, TE strips were

lyophilized after mechanical testing (n=4-5 per group) and used for biochemical assays.

The total amount of DNA was determined as an indicator of cell number, the amount of

hydroxyproline as an indicator for collagen content, and the amount of sulfated


25

2

glycosaminoglycans (sGAG). Measurements were averaged per group. Lyophilized tissue

samples were weighted and digested in papain solution (100 mM phosphate buffer

(pH=6.5), 5 mM L-cysteine, 5 mM ethylene-di-amine-tetra-acetic acid (EDTA), and 125-

140 μg papain per ml, all from Sigma) at 60°C for 16 hours. After centrifuging the samples,

the digest supernatant was collected and used for the DNA, sGAG and collagen assays.

The amount of DNA in the TE strips was determined using the Hoechst dye method [162]

and a standard curve prepared of calf thymus DNA (Sigma). Using the assumption that all

cells contain 6.5 pg of DNA [163], the amount of cells per TE construct was calculated.

sGAG content was determined with a modification of the protocol described by Farndale

et al. [164]. In short, 40 μl of diluted sample was pipetted into a 96-well plate in duplo

followed by addition of 150 μl di-methyl-methylene blue per well. Absorbance was

measured at 540 and 595 nm and extracted from each other. Subsequently, the amount

of sGAGs in the TE strips was determined from a reference curve prepared from shark

cartilage chondroitin sulfate (Sigma). Collagen content was determined by an assay as

described by Huszar et al [165], and a standard curve was prepared from trans-4-

hydroxyproline (Sigma).

2.3.6 Mechanical testing

After 4 weeks of culture, the mechanical properties of the TE strips (n=4-5 per group) were

assessed by uniaxial tensile tests in longitudinal direction with a uniaxial tensile stage

(Kammrath &Weis, Dortmund, Germany) equipped with a 20N load cell. Mechanical test

data was averaged per group. Thickness of the strips was determined from representative

histology sections. Samples were measured at three spots and mean thickness was used.

Standard deviation of the measurements ranged 1.5-10%. Stress-strain curves were

obtained and as a measure for tissue strength, the ultimate tensile strength (UTS) was

defined as the peak stress value. The elasticity modulus (E-modulus) was determined as

the slope of the linear (end) part of the curve, as a measure for tissue stiffness.

2.3.7 Histology

To analyze tissue formation qualitatively, TE strips were processed for histology (n=1 per

group). Representative tissue samples were embedded in tissue freezing medium (Tissue

Tek, Sakura, Torrance, USA) and cryosections of 10 μm were cut. The sections were

formalin-fixed and studied by Masson Trichrome (MT) staining (MTC kit, Sigma, Venlo,

Netherlands) for collagen deposition and by Picrosirius Red (PR) staining to assess the

maturity of the collagen matrix [166]. The MT staining was analyzed using light microscopy

and the PR staining by polarized light microscopy (Axio Observer, Zeiss, Göttingen,

Germany). In this study, maturity of the collagen fibers was assessed by amount and

density of the collagen fibers visible with polarized light microscopy. Mature fibers with a

high density are colored orange/red, while immature or less dense fibers are green.

Chapter 2

26

Cell phenotype within the TE strips was analyzed by immunofluorescence. After acetone

fixation for 10 minutes, the sections were incubated with 5% bovine serum albumin (BSA)

in PBS for 30 minutes at room temperature. After blocking, the sections were incubated

with a primary antibody overnight at 4°C. Antibodies used were mouse anti α-smooth

muscle actin (αSMA) to stain smooth muscle cells and myofibroblasts (a2547, clone 1A4,

Sigma, 1:400), mouse anti-smoothelin to stain contractile smooth muscle cells (Clone R4A,

kindly provided by GJ van Eys from the University Maastricht, 1:4) or rabbit anti S100A4,

which belongs to the S100 superfamily of cytoplasmic calcium-binding proteins, to stain

fibroblasts and myofibroblasts (ab27957, Abcam, 1:200). After primary antibody

incubation, the sections were washed with PBS and incubated with Alexa 488 labeled

secondary antibodies (Sigma and Molecular probes, 1:300) to visualize the specific

stainings and DAPI (Sigma, 1:500) to stain all cell nuclei for 30 minutes at room

temperature. After staining, sections were mounted with Mowiol 4-88 (Calbiochem, San

Diego, USA) and visualized by fluorescent microscopy (Axiovert 200M, Zeiss, Göttingen,

Germany).

2.3.8 Statistical analyses

All data are presented as mean ± standard error of the mean. Data of all experiments were

normalized to human passage 3 PGA-P4HB strips in each experiment to be able to

compare experiments and perform statistical analyses. Pearson correlation coefficients

were calculated to determine correlations between tissue parameters and cell passage

numbers for both species and scaffold groups. Unpaired t-tests were used to compare the

tissue properties between the scaffold materials within one cell passage and species, and

to compare the tissue properties between species, within the same scaffold material and

cell passage number. Statistics were performed using GRAPHPAD Prism (version 5.04) and

differences were considered significant for p-values <0.05.

2.4 Results

2.4.1 Compaction after 4 weeks

The remaining width of the strips of all groups after 4 weeks of culture is shown in Figure

2.1 A. A remaining width of strips of 100% is the initial width of the strips and represents

no compaction. The tissues based on PCL scaffold, and PCL and PGA-P4HB control strips,

showed compaction of 0-10%. The tissues based on PGA-P4HB scaffold resulted in

significant more compaction of around 50% after 4 weeks (p<0.001).

In ovine strips, no significant correlation between passage number and both types of

scaffold was found. A negative correlation was found between human cell passage

numbers and PGA-P4HB strips (p<0.01), while there was a positive correlation between

the human cell passage numbers and PCL strips (p<0.05). This indicates that passage


27

2

number and species did not consistently influence compaction. TEHV based on PGA-P4HB

scaffolds show severe compaction and retraction of the leaflets after 4 weeks culture in

both species, while no compaction or retraction was observed in the PCL based valves

(Figure 2.1 B-E), confirming the results as found in the engineered strips.

Figure 2.1 Compaction of strips after 4 weeks of culturing. Initial width of strips was set at 100% (dotted

line) (A). PGA-P4HB showed around 50% compaction of the strips, while the use of PCL strips demonstrated

reduced compaction as the final reduction in width was 0-10% only. ** indicates the difference between

the scaffold materials with a p-value<0.001, while # and ## denote significant differences of p<0.05 and

p<0.001 compared to human tissues respectively. Negative or positive Pearson r correlations between the

cell passage numbers are presented by arrows combined with their p-values. Species and cell passage

number did not consistently influence compaction of the TE strips. Top view photos of a human PGA-P4HB

(B), human PCL (C), ovine PGA-P4HB (D) and ovine PCL (E) TEHV after 4 weeks of culture. Valves based on

PGA-P4HB scaffold resulted in severe retraction of the leaflets after 4 weeks, while PCL valves did not show

this. These results were consistent for both human and ovine cells.


Normalized collagen and sGAG per DNA of all groups are presented in Figure 2.2.

Significant negative correlations between cell passage numbers and collagen amount per

DNA, were found in both human and ovine tissues of both scaffold materials (p<0.001),

demonstrating that increasing passage number, resulted in decreased collagen per DNA.

Low amount of collagen per DNA was detectable in ovine PCL p7 strips. In general ovine

tissue strips demonstrated an increased amount of collagen when compared to human

(p<0.001). Collagen content per DNA of both human and ovine p7 cells seeded on PCL

scaffolds was decreased, compared to human and ovine cells that were seeded on PGA-

P4HB scaffolds (p<0.05 for human cells and p<0.001 for ovine cells). Although we showed

that collagen and sGAG per DNA was increased with decreasing passage number, no

differences in compaction of the tissues could be observed.

Biochemical parameters are related, as observed by correlation matrices, showing that

collagen per DNA was increased when sGAG per DNA was increased. Overall, the amount

Chapter 2

28

of sGAG per DNA decreased with increasing cell passage (p<0.05 for ovine PGA-P4HB

strips and p<0.001 for human PCL strips) although this effect was less pronounced as seen

for to collagen per DNA. Except ovine p7 PCL strips, ovine cells resulted in a higher amount

of sGAG per DNA compared to human cells (p<0.05 for ovine p3 PCL strips and p<0.001

for all other ovine strips). No consistent differences in sGAG content by the cells were

observed due to different scaffold materials.

Figure 2.2 Collagen per DNA (A) and sGAG per DNA (B). # and ## denote significant differences compared to human tissues, while * and ** denotes significances of differences between scaffold materials with p<0.05 and p<0.001. Pearson r correlations between the cell passage numbers are presented by arrows combined with their p-values. Collagen per DNA is decreasing with increasing passage number in both human and ovine tissues and both scaffold materials (A). sGAG per DNA show the same trends although less distinct (B). Scaffold does not influence the amount of formed collagen and sGAG, while per DNA, more collagen and sGAG are formed within ovine tissues compared to human tissues.


In Figure 2.3A and 2.3B, averaged stress strain curves of the human and ovine p3 strips,

which are representative for the other passage numbers, and the PGA-P4HB and PCL

control strips are presented. Bare PCL strips were able to bear higher stresses compared

to bare PGA-P4HB strips, which is due to the differences in degradation time of both

scaffold materials. The PGA-P4HB cultured tissues of both human and ovine cells showed

typical non-linear mechanical behavior representing tissue behavior. When PCL scaffold

was used, human tissues showed linear mechanical behavior, while the ovine tissues were

following the curve of the control PCL strips. Thus, PCL scaffolds are still influencing the

mechanical properties of the engineered tissues after 4 weeks of culture, while PGA-P4HB

scaffolds do not.


29

2

Figure 2.3 Mechanical data of engineered strips. Averaged stress strain curves of human (A) and ovine (B)

p3 strips are given as mean ± SEM. PGA-P4HB based tissues demonstrate non-linear curves in both human

and ovine strips, representing tissue behavior. The stress strain curve of human PCL strips is linear, while

ovine strips follow the curve of the control scaffolds. Control PCL scaffolds are still influencing mechanical

properties after 4 weeks of culture, while PGA-P4HB scaffolds are not. Tissue stiffness (C) and strength (D)

are increasing with decreasing passage number. # and ## denote significant differences compared to human

tissues, while * and ** denotes significances of differences between scaffold materials with p<0.05 and

p<0.001. Pearson r correlations between the cell passage numbers are presented by arrows combined with

their p-values. In human samples, highest values are obtained in PCL strips, while in ovine this is observed

in PGA-P4HB scaffold strips.

With a decrease of cell passage numbers, the parameters stiffness and strength were

increasing in both species and scaffold materials, as significant negative correlations were

observed between increasing cell passage numbers and both the stiffness (p<0.05 for

human PGA-P4HB strips and p<0.001 for human PCL and ovine PGA-P4HB strips) and

strength (p<0.05 for human PGA-P4HB and ovine PCL strips and p<0.001 for human PCL

and ovine PGA-P4HB strips), in human and ovine tissues based on both scaffold materials

(Figure 2.3C and 2.3D). In human tissue samples, stiffness was higher in PCL samples

compared to PGA-P4HB samples (p<0.05 in p3 and p7 tissues), while in ovine tissue

samples a higher stiffness was obtained in tissues based on PGA-P4HB scaffolds compared

to PCL scaffolds (p<0.05). Furthermore, tissue strength was increased in human PCL

samples of all passage numbers and ovine PCL samples of passage 5 and 7, compared to

PGA-P4HB tissue samples (p<0.05) which probably is due to the influence of the PCL

scaffold that is not yet degraded. When PCL scaffold was used, the values of the

Chapter 2

30

mechanical properties of the ovine tissues were equally or just slightly increased

compared to the PCL control strips, while the values of the human tissues were higher

compared to the control strips (data not shown). This indicated that the newly formed

tissues by ovine cells were not of the same quality as their human equivalents, as the

added value of tissue to the mechanical properties of the ovine strips was relatively low.

Correlation matrices demonstrated that mechanical parameters are related to each other,

resulting in increased tissue strength when tissue stiffness obtained higher values, while

mechanical parameters were not related to matrix properties of the tissues.

2.4.4 Histology

Histology of the TE strips revealed cellular tissue with dense surface layers. Picrosirius Red

and Masson Trichrome stainings (Figures 2.4 and 2.5), showed collagen fibers in strips of

all groups after 4 weeks of culture. A higher amount of red fibers was seen in most tissues

with cells of a low passage number (Figure 2.4). This indicated that tissues based on a low

cell passage number resulted in more mature collagen fiber formation. Histology

furthermore indicated that the total amount of collagen fibers was decreasing with

increasing passage numbers in both PGA-P4HB and PCL strips (Figure 2.5). Ovine PGA-

P4HB tissues showed a higher amount of collagen compared to the human tissues.

However, ovine PCL based tissues showed little amount of collagen compared to human

PCL based tissues. The total amount of collagen was higher in PGA-P4HB strips compared

to PCL strips, which can be explained by triple the amount of cells seeded onto the PGA-

P4HB strips compared to PCL strips, due to differences in thickness of the scaffold

materials. Immunofluorescent stainings indicated no differences in cell phenotype with

cell passage number, scaffold material or species, as tissues of all groups contained cells

that were αSMA and S100A4 positive, and smoothelin negative (Figure 2.6), indicative for

synthetic myofibroblasts. Cells in all strips were distributed homogenously throughout the

strips as shown by cell nuclear staining (DAPI).


31

2

Figure 2.4 Picrosirius Red stained sections of human PGA-P4HB (A-C), PCL (D-F), ovine PGA-P4HB (G-I) and PCL (J-L) visualized by polarized light microscopy. Maturity of collagen fibers is visualized as green (immature) and orange/red (mature). Most red fibers are visualized in tissues based on cells with a low passage number, indicating that maturity of collagen fibers after 4 weeks of culture is decreasing with increasing passage number. The white scale bars represent 200 µm. Vacuoles within the scaffolds are cutting artifacts due to scaffolds remnants, and grey parts in the PGA-P4HB groups are P4HB remnants.

P3 P5 P7

Ovi

ne

PC

L O

vin

e P

GA

-P4

HB

H

um

an P

GA

-P4

HB

H

um

an

PC

L

Chapter 2

32

Figure 2.5 Masson Trichrome staining of human PGA-P4HB (A-C), PCL (D-F), ovine PGA-P4HB (G-I) and PCL (J-L) sections. The blue scale bars represent 200 µm. Collagen is shown in blue and red represents cytoplasm and muscle tissue. Vacuoles within the scaffolds are cutting artifacts due to scaffolds remnants. The total amount of collagen fibers seem to decrease with increasing passage number in both scaffold materials. Ovine PGA-P4HB strips show more collagen compared to human strips, while in PCL strips most collagen is visualized in human samples.

Figure 2.6 Representative photos of immunofluorescent stainings of the αSMA (A), S100A4 (B) and the Smoothelin (C) cell markers, with the white scale bars representing 200 µm. In green the protein of interest is colored, in blue DAPI is visible to stain cell nuclei. All stained tissues contain cells that were αSMA and S100A4 positive and smoothelin negative. This indicates that passage number, scaffold material and species are not influencing cell phenotypes. Vacuoles within the scaffolds are cutting artifacts due to scaffolds remnants.

P3

Ovi

ne

PC

L O

vin

e P

GA

-P4

HB

H

um

an P

CL

Hu

man

PG

A-P

4H

B

P5 P7


33

2

2.5 Discussion

Compaction and retraction of heart valve leaflets in vitro, resulting in regurgitation in vivo,

is a common problem in TEHV that are based on rapid degrading PGA-P4HB scaffolds.

Therefore, alternative approaches to overcome compaction and retraction of TEHV are

needed to meet in vivo demands. This study has focused on the effect of two alternative

approaches: 1) the use of a slow degrading PCL scaffold and 2) the use of lower passage

vascular cells. Compaction, tissue formation, cell phenotype and mechanical properties of

both human and ovine tissues were investigated.

2.5.1 Differences due to vascular cell expansion times

In this study, we demonstrated that reduced in vitro expansion time of vascular cells

resulted in improved tissue amount as sGAG per DNA, collagen per DNA, tissue strength

and stiffness were increased with decreasing passage number. A comparison of the net

amounts of collagen and sGAG could not be made, as different amounts of cells were

seeded, due to differences in scaffold thickness. Therefore, collagen and sGAG were

normalized to DNA. A 2D study of ovine jugular vein derived cells showed that sGAG

content was highest in low passage cells [167]. Although cells in 2D may act differently

compared to cells in 3D, our data also indicated that cells with an increasing passage

number became less synthetic, as collagen and sGAG content was decreased by cells of a

higher passage number. Some in vitro studies showed that the vascular contractile

smooth muscle cell marker smoothelin, disappeared within a few days of in vitro

expansion, and cells differentiated into synthetic, tissue producing cells [168], while

others observed this only after the 9-11th passage [169, 170]. All our human and ovine

cells have been differentiated into the synthetic phenotype, as no change of phenotype

could be observed in this study due to cell passage number, and all tissue sections showed

αSMA and S100A4 positive, and smoothelin negative cells indicating activated, synthetic

myofibroblasts. Cell phenotype of our samples and amount of tissue were not related as

no change in cell phenotype could be observed, while it was shown that the amount of

tissue increased with decreasing passage numbers.

The tissue stiffness of strips was obtained from the linear end part of the stress strain

curves and represents the end stiffness of our tissues. Increase of tissue stiffness, was

seen in strips based on a decreased cell passage number. The increase in end stiffness of

our tissues resulted in stronger tissues, although, native leaflets still do show much higher

values of stiffness compared to our tissues; 15.6 ± 6.4 MPa in the circumferential direction

and 2.0 ± 1.5 MPa in the radial direction [171]. Native valves are also more flexible

compared to our engineered strips when comparing the physiological relevant stiffness.

The opening and closing functions of the heart valves are controlled by pressure

differences. As the native valves are more flexible compared to their engineered

counterparts, a lower pressure is needed for opening the valves.

Chapter 2

34

Histology of the PGA-P4HB samples confirmed the biochemical results of collagen content

per DNA, as higher amounts of collagen were observed in the ovine PGA-P4HB tissues

compared to the human tissues. This increased amount of collagen in ovine PGA-P4HB

based tissues is not only explained by increased synthetic ovine cells, but also by a higher

proliferation rate of these cells compared to human cells when seeded on PGA-P4HB

scaffolds (proliferation data not shown). However, ovine PCL based tissues show little

collagen in the histology slides compared to human PCL based tissues, while the

biochemical data showed an increased amount of collagen per DNA in ovine tissue

compared to human. This can be explained by the proliferation rate of ovine and human

cells in PCL scaffolds. As human cells showed a higher proliferation rate when seeded onto

PCL scaffolds (data not shown) and, therefore, an increased amount of total DNA per strip

in PCL scaffolds compared to ovine cells, the amount of collagen per DNA is lower in

human, while the total amount of collagen per strip might be higher due to the presence

of more collagen producing cells. More research is needed to investigate why differences

in proliferation rates of ovine and human cells are present when different types of

scaffolds are used.

Mechanical results also correlated with the histological findings. Strips that showed more,

and increased maturity of collagen fibers, also resulted in an increased tissue stiffness and

strength. This is in line with previous findings, where a dominant role for collagen maturity

by cross-linking of the collagen over collagen content was found with respect to

mechanical properties of the tissues [171].

Remarkable is that ovine p7 PCL strips resulted in only few cells present after 4 weeks.

Collagen content of these cells was also low resulting in weak strips as observed in the

tensile tests. We hypothesize that this might be due to the combination of several factors.

One might be the use of a low amount of serum (0.1% in ovine 3D medium). This could

have resulted in non-synthetic and non-dividing cells. In combination with the high

passage number, which also showed to result in less activated or synthetic cells, this could

have been the reason for the low amount of cells present after 4 weeks and reduced

amount of collagen. Furthermore, the use of PCL scaffold is likely to have influenced the

amount of collagen, as ovine p7 cells seeded on PGA-P4HB scaffolds, did show higher

amounts of collagen. We hypothesized that the use of PCL scaffold with ovine cells,

resulted in non-synthetic cells, as the mechanical integrity of this scaffold was present for

a longer time span, resulting in no urgent need for the cells to create tissue. However,

culturing TEHV with ovine p7 cells did result in proper tissue formation. This might be

explained by different culture protocols of engineered strips and TEHV. TEHV undergo

mechanical loading in a bioreactor during culture, while strips are cultured statically.

Furthermore, interspecies differences might have played a role, as cells of a different

sheep were used to culture the TEHV.


35

2

Concerns might rise about the clinical applicability of using cells with a low passage

number, mainly in children, as a relatively large number of cells need to be obtained.

However, in the case of children fewer cells are needed to be able to produce a TEHV

compared to adults, as the annulus of the pulmonary valve in children is 10-17 mm, while

this is around 25 mm in adults. The size of the leaflets in young patients is also smaller.

Furthermore, when PCL based TEHV are produced instead of PGA-P4HB based TEHV,

fewer cells are needed due to differences in scaffold thickness. To produce a PCL based

TEHV scaffold for adults, 20 x 106 cells are needed, while this would be 2-10 x 106 cells is

case of children. These amounts of cells can be obtained by the outgrowth method as the

saphenous vein segments need to be a centimeter only. In conclusion, cells from a lower

passage number demonstrated to increase the amount of tissue formation and tissue

strength, without influencing cell phenotype. Despite the improved tissue formation,

compaction of the tissues was not influenced by a lower cell passage number.

2.5.2 PGA-P4HB versus PCL scaffold

In this study, we demonstrated that human and ovine tissues cultured for 4 weeks using

PCL scaffold strips showed almost no compaction (0-10%), while PGA-P4HB based tissues

showed compaction up to 50%. Furthermore, we showed that PGA-P4HB based TEHV

resulted in severe retraction of the leaflets in both species, while this was not seen in the

PCL based TEHV. This proves that PCL is a promising scaffold material to reduce

compaction and retraction in TEHV. Dijkman et al described another approach to prevent

compaction and retraction of PGA-P4HB based TEHV [43]. Trileaflet heart valve of PGA-

P4HB scaffolds were seeded with ovine myofibroblasts and subsequently decellularized

to prevent retraction. Decellularization represented to be a powerful tool to reduce tissue

retraction, as it was shown that cell-induced retraction accounted for 85% of total tissue

retraction. Residual matrix stresses are known to still account for 15% of the total

retraction [74]. These residual matrix stresses minimized the coaptation area in the study

of Dijkman et al. and it has to be elucidated in future studies whether this will influence

in vivo valve behaviour. We believe that by using a slow degrading scaffold, retraction can

be even more effectively reduced by resisting residual matrix stresses, while maintaining

tissue viability.

Results of the mechanical tests demonstrated that in PCL strips the mechanical properties

were not only determined by the formed tissue, but also by the remaining PCL scaffold,

as it was not yet degraded. PGA-P4HB is known to start to degrade after 2 weeks, and,

therefore, was not influencing the mechanical properties of the tissues. As amounts of

sGAG and collagen per DNA were not influenced by the scaffold materials, the increased

tissue strength of the human PCL strips compared to the PGA-P4HB strips are likely due

to the remaining PCL scaffold. Ovine strips did not show the same results, which might be

due to the low amount of DNA and, therefore, a lower amount of total tissue, in ovine PCL

Chapter 2

36

strips. Mechanical properties of the ovine PCL strips were mainly influenced by the

remaining scaffold and not by the formed tissue, while in PGA-P4HB strips the measured

mechanical properties represented tissue only. Furthermore, ovine tissues based on PCL

scaffold did not influence mechanical properties as much as compared to human PCL

tissues, as tissue strength and stiffness values were equally or just slightly increased

compared to the PCL control strips. This indicated that the newly formed tissues based on

ovine cells were not of the same quality as their human equivalents.

Differences in scaffold thickness could possibly have resulted in differences in tissue

formation, due to variation in nutrient and oxygen levels within the tissues. This is mainly

seen in ovine strips as human strips show more homogeneously distributed tissue. Our

ovine strips possess a denser layer of collagen and cells on the surface in both scaffold

material. However, cells were not only present at the surface layer, but also distributed

throughout the center of both scaffold materials. Not only the cells at the surface layer,

but also the cells in the center produced collagen and expressed the synthetic smooth

muscle cell markers, as visualized by histology. Furthermore, biochemical assays

demonstrated no influence of the scaffold materials on the collagen and sGAG formation

per DNA, and differences in mechanical properties of the tissues are most likely due to PCL

scaffold remnants instead of differences between material thicknesses. Directly after

seeding, the high porosity of the scaffold strips allowed oxygen and nutrient supply to the

cells that were situated on the scaffold fibers in the middle part of the strip. When tissue

was produced, porosity decreased and oxygen and nutrient supply might have been

decreased resulting in the formation of surface layers.

Native human heart valve leaflets are avascular as they are thin enough to receive

nutrients and oxygen through diffusion and hemodynamic convection [18]. As PCL

scaffolds are 300 µm, we do not expect problems when placing PCL TEHV in vivo. TEHV

based on PGA-P4HB did show increased thickness in the ovine model [149], which might

lead to reduced oxygen and nutrient supply to the cells in the center. This problem might

be less pronounced in human as these tissues are also compacting in the vertical direction,

and therefore decreasing in thickness.

In conclusion, the use of PCL scaffold seems to be an alternative scaffold material for the

culture of human TEHV to reduce compaction, while further optimization is needed when

ovine cells are used to ensure proper tissue formation.

2.5.3 Interspecies differences

Tissue properties were different between species. In our study, ovine cells presented to

be more synthetic compared to human cells as they contained more sGAG and collagen

per DNA, while a study by van Geemen et al, demonstrated the opposite effect [48]. Van

Geemen showed that human passage 7 cells contained double the amount of sGAG per


37

2

DNA (4.8 ± 0.8 µg/µg DNA in ovine and 8.2 ± 1.4 µg/µg DNA in human cells) and five times

the amount of collagen per DNA (1.1 ± 0.3 µg/µg DNA in ovine and 5.9 ± 2.5 µg/µg DNA in

human cells) compared to ovine passage 7 cells. Tissues based on passage 7 cells in our

experiments obtained values for sGAG per DNA of 6.5 ± 0.2 µg/µg ovine DNA and 5.5 ±

0.3 µg/µg human DNA. Collagen per DNA was 3.2 ± 0.1 µg/µg DNA, and 3.7 ± 0.3 µg/µg

DNA, in ovine and human respectively. This suggests that ovine cells in our study were

more synthetic or less proliferative, which might be due to the amount of serum used in

the culture medium. Van Geemen used 2.5% of lamb serum, while in this study 0.1%

serum was used only, as an in vitro TEHV study by Dijkman demonstrated more

homogeneous tissue formation throughout the wall and leaflets when 0.1% lamb serum

was used [172]. A review by Mol et al described that the outcome of ovine TEHV was

dramatically different from their human equivalents when using the same culture

conditions, and lower amounts of serum resulted in tissue outcome comparable to human

[173]. This shows the difficulties in the translation step from animal studies towards the

clinic and vice versa. Furthermore, previous studies showed that not only interspecies,

but also intraspecies variations of tissue properties are large [48, 171, 174]. Within this

study we investigated the tissue properties of the strips seeded with cells of one sheep

and one patient only. While it would be preferred to have more data on several human

and ovine cell sources, we assume that within species the effects of e.g. cell passage

number are comparable. Furthermore, the first goal of this study was to compare

different types of scaffold to prevent compaction. This was investigated on cells of two

species (human and ovine) and different cell passage numbers of those species. While two

species and cell passage numbers were used and differences in terms of tissue production

were observed between these species and cells passage numbers, the outcome of

compaction was similar in all research groups. This indicates that the influence of the

scaffold type is larger as compared to the influence of the tissue production of several cell

sources, in terms of compaction.

A limitation of our study is that the ovine cells originated from a young, healthy sheep,

while the human vascular derived cells were obtained from an older person that

underwent bypass surgery. This might have influenced the outcome of the tissue

properties as not only cell passage number, but also patient age may have an effect on

the cell functioning, doubling time and ability of tissue production in different cell types

[160, 175, 176].

In conclusion, differences in absolute values between ovine and human samples were

seen within this experiment, although the general effects of reducing cell passage

numbers and the use of PCL scaffold on compaction and the amount of tissue formation

were comparable.

Chapter 2

38

2.6 Conclusion

This study showed that PCL scaffolds may serve as alternative scaffold material for human

TEHV with minimal compaction and without compromising on tissue composition and

properties, while further optimization of ovine TEHV based on PCL scaffold is needed to

not only ensure reduced compaction but also strong tissues of a high quality. Cells from

lower passages demonstrated to improve tissue formation, without influencing

compaction and cell phenotype. In addition, reducing cell expansion will result in faster

generation of TEHV, providing a more rapid treatment to patients.

Acknowledgements

This work was supported by a grant from the Dutch government to the Netherlands

Institute for Regenerative Medicine (NIRM, grant No. FES0908). The authors wish to thank

Tom Lavrijsen, Leonie Grootzwagers and Anita van de Loo for their help with the

mechanical tests and culturing the TEHV. Furthermore Marc Simonet is acknowledged for

the production of PCL scaffolds. The smoothelin antibody was kindly provided by Dr. GJ

Van Eys, department of molecular genetics, cardiovascular research institute Maastricht,

University Maastricht.

39

Superior tissue evolution in slow-

degrading scaffolds for valvular

tissue engineering

3

M. Brugmans

R. Soekhradj-Soechit

D. van Geemen

M. Cox

C. Bouten

F. Baaijens

A. Driessen-Mol

Submitted

Chapter 3

40

3.1 Abstract

Synthetic polymers are widely used to fabricate porous scaffolds for the regeneration of

cardiovascular tissues. To ensure mechanical integrity after implantation, a balance

between the rate of scaffold resorption and tissue formation is of high importance. In

vivo, a higher rate of tissue formation is expected in fast-resorbing materials compared to

slow-resorbing materials, as a result of highly synthetic cells, which aim to compensate

for the fast loss of mechanical integrity of the scaffold by deposition of newly formed

collagen fibers. Here, we studied the effect of fast- (PGA-P4HB) and slow-degrading (PCL)

synthetic scaffolds on tissue amount, composition, and mechanical characteristics in time

in vitro, and compared these engineered values with values for native human heart valves.

Electrospun porous PGA-P4HB and PCL scaffolds were either kept unseeded in culture or

were seeded with human vascular-derived cells. Tissue formation, ECM composition,

remaining scaffold weight, tissue to scaffold weight ratio, and mechanical properties were

analyzed weekly up to 6 weeks. Unseeded PCL scaffolds remained stable in weight during

the 6-week culture, while PGA-P4HB scaffolds degraded rapidly. When seeded with cells,

both scaffold types demonstrated increasing amounts of tissue with time, which was

more pronounced for PGA-P4HB-based tissues during the first two weeks due to highly

synthetic cells, however PCL-based tissues resulted in the highest amount of tissue after

6 weeks. This study is the first to provide insight into the tissue to scaffold weight ratio,

therewith allowing for a fair comparison between engineered tissues cultured on scaffolds

with different degradation rates, as well as to native heart valve tissues. Although the

absolute amount of ECM components differed between the engineered tissues, the ratio

between ECM components was similar after 6 weeks. PCL-based tissues maintained their

3D shape during culture, while the deformed PGA-P4HB-based tissues showed

appositional growth with culture time. After 6 weeks, PCL-based engineered tissues

showed amounts of cells, collagen, and glycosaminoglycans that were comparable to

human native heart valve leaflets, while engineered values were lower in the PGA-P4HB-

based tissues. Although increasing in time, the amounts of collagen crosslinks were still

below native values in all engineered tissues. In conclusion, this study indicates that slow-

degrading scaffold materials are favored over fast-degrading materials in order to create

organized ECM-rich tissues in vitro, which keep their 3D structure before implantation.

Superior tissue evolution in slow-degrading scaffolds

41

3

3.2 Introduction

Bioresorbable synthetic polymers are used extensively in the field of cardiovascular tissue

engineering to fabricate three-dimensional porous scaffolds, aiming for the regeneration

of different types of tissues, such as heart valves and blood vessels [72, 177]. The classical

tissue engineering paradigm to develop tissue replacements is the in vitro tissue

engineering approach, where cells are seeded into synthetic bioresorbable porous

scaffolds. During subsequent culture, tissue will be produced by the cells and after culture

the construct can be implanted as a living autologous replacement. Alternatively, the

engineered tissue can be decellularized after culture, to create allogenic off-the-shelf

replacements that are rapidly repopulated to function as living replacement [43, 73, 75,

178].

As we focus on cardiovascular applications, we use human primary vascular-derived cells

to grow tissue in fast-degrading PGA-P4H-based scaffolds, or slower-degrading PCL-based

scaffolds. Both scaffold types have previously shown excellent results in terms of

biocompatibility, processing ability, and cell infiltration [43, 47, 58, 75, 179, 180]. In vivo,

scaffolds made of these materials will be fully resorbed by the body, ultimately resulting

in a living implant that is able to adapt and remodel. Different research groups have

studied the resorption of these scaffolds, mainly in vivo. It appeared that the site of

implantation, presence of enzymes, molecular weight of the material, and scaffold

porosity all affect resorption rates in vivo [181]. Reported complete resorption times of

PGA vary from 1.5 months [68] to 4-6 months [182]. For electrospun PCL scaffolds,

resorption takes much longer and this type of scaffold is reported to be completely

resorbed in vivo after at least 2 years [183]. Obviously, scaffolds with slow- and fast-

resorption rates will contribute to the mechanical integrity of the tissue differently with

time, both in vivo and in vitro. We hypothesize that cells in fast-resorbing scaffold

materials will comprehend increased rates of tissue production, in order to compensate

the loss of mechanical integrity by formation of collagen fibers, compared to cells in slow-

resorbing scaffold materials where mechanical integrity is maintained for a longer period

of time. How tissue composition is changing during in vitro culture, and how this affects

mechanical integrity due to different degradation properties of the scaffold materials, was

never fully assessed. In an attempt to balance scaffold degradation, tissue stability, and

mechanical integrity for in vitro cardiovascular tissue engineering, we determined the

weight ratio between scaffold and tissue weekly, during a 6-week culture period, using

both a slow- and a fast-degrading scaffold. In addition, we analyzed absolute and relative

amounts of ECM and mechanical properties of the constructs with time.

Previously, attempts were undertaken to compare tissue composition of engineered

constructs, cultured on different types of bioresorbable scaffolds, to native cardiovascular

tissues [68, 80, 93, 184]. Results from conventional tissue composition assays express

tissue composition relative to the dry weight of the sample. Often the weight of the

Chapter 3

42

remaining scaffold is integrated into the weight that is used to compare tissue

composition of engineered tissues with native tissues [182, 185-187]. This results in an

overall overestimation of the actual formed tissue weight and, therewith, an

underestimation of the amount of ECM components per mg formed tissue. To make a fair

comparison between engineered and native tissues, the obtained data should be

corrected for the remaining scaffold weight. During ageing, cardiovascular tissues are

known to change in terms of ECM composition. We compared our engineered

cardiovascular tissues to native valvular human data of several age groups as determined

by van Geemen et al.[188], to assess similarity of the engineered tissues to native tissues.


3.3.1 Cell culture

Human vascular-derived cells were harvested from segments of a Vena Saphena Magna

from a patient that underwent bypass surgery, and was obtained according to the Dutch

guidelines for secondary used materials. Cells were obtained using the outgrowth method

and cultured using standard culture methods in a humidified atmosphere containing 5%

CO2 at 37°C, as described previously [47]. Isolation and expansion medium consisted of

advanced Dulbecco’s Modified Eagle Medium (DMEM; Invitrogen, Breda, Netherlands),

supplemented with 1% GlutaMax (Gibco), 1% Penicillin/Streptomycin (P/S, Lonza, Basel,

Switzerland), and 10% Fetal Bovine Serum (Greiner Bio-one, Alphen a/d Rijn,

Netherlands).


Rectangular strips (25x5 mm) were cut out of PGA meshes (PGA; specific gravity, 70

mg/cm3; Cellon, Bascharage, Luxemburg) and conventionally electrospun PCL meshes,

both with a thickness of 1 mm and comparable fiber diameter. PGA scaffolds were

additionally coated with poly-4-hydroxybutyrate (P4HB, received via a collaboration with

Prof. Hoerstrup of the University Hospital Zurich) to provide structural integrity to the

mesh. The outer 3-4 mm of both PGA and PCL scaffold strips were attached onto stainless

steel rings (RVS Paleis, Geleen, Netherlands) using 15% polyurethane-tetrahydrofuran

(PU, DSM, Geleen, Netherlands) glue, leaving an 18*5 mm area for cell seeding. The

solvent was allowed to evaporate overnight in a vacuum oven. The rings with the scaffold

strips were placed in 6-well plates and sterilization was achieved by immersion in an

antibiotic/anti fungi solution, consisting of 10% Penicillin/Streptomycin (P/S; Lonza) and

50 µg/ml Fungin in sterile Phosphate Buffered Saline (PBS) (Sigma, Venlo, Netherlands)

for 30 minutes on a shaker at 37˚C. Subsequently, the antibiotics/anti fungi solution was

removed and 70% ethanol was added for 15 minutes. The ethanol step was repeated and,

thereafter, the strips were washed twice in PBS. To facilitate cell attachment, the scaffolds


43

3

were incubated overnight with tissue engineering (TE) medium, consisting of expansion

medium supplemented with 0.25 mg/ml L-ascorbic acid 2-phosphate (Sigma).

3.3.3 Experimental design

Scaffold strips of both materials (n=58) were either kept unseeded in culture (n=4 per

week) or were seeded with cells (n=4-5 per week). Passage 7 cells were used and seeded

onto both PGA-P4HB and PCL scaffolds, with a seeding density of 2.0 x 106 per cm3 using

fibrin as a cell carrier [161]. In short, cells were suspended in TE medium containing

thrombin (10 U/ml, Sigma). This cell suspension was mixed with an equal volume of TE

medium containing fibrinogen (10 mg/ml, Sigma) and dripped onto one side of the

scaffolds. After seeding, the constructs were placed in an incubator at 37°C for 30 minutes,

to allow polymerization of the fibrin gel. Thereafter, 6 ml of TE medium was added to each

scaffold. The constructs were cultured for up to 6 weeks and TE medium was changed

twice a week. After 1, 2, 3, 4, 5, and 6 weeks, seeded strips (n=4-5 per week) were

sacrificed. One strip was used for histology and the remaining strips were used for

mechanical testing followed by biochemical assays. At week 0, 1, 2, 3, 4, 5, and 6,

unseeded strips (n=4 per week) were sacrificed. These strips were used for mechanical

testing only.


For the quantification of tissue formation during culture, engineered constructs were

lyophilized after mechanical testing (n=3-4 per group) and used for biochemical assays.

The total amount of DNA was determined as an indicator of cell number, the amount of

hydroxyproline (hyp) as an indicator for collagen content, and the amount of sulfated

glycosaminoglycans (sGAG) was measured. Measurements were averaged per group.

Lyophilized constructs were weighed and digested in papain solution (100 mM phosphate

buffer (pH=6.5), 5 mM L-cysteine, 5 mM ethylene-di-amine-tetra-acetic acid (EDTA), and

125-140 μg papain per ml, all from Sigma) at 50°C for 16 hours. To compare DNA, sGAG

and collagen within engineered constructs in time and with values found in native tissue,

the weight of the engineered tissues without scaffold needs to be calculated. To obtain

these values, the weight of the remaining scaffold of the unseeded strips of equal time

points was subtracted from the weight of the seeded strips. The digest supernatant was

collected and used for the DNA, sGAG and collagen assays. The amount of DNA in the

constructs was determined using the Hoechst dye method [162] and a standard curve

prepared of calf thymus DNA (Sigma). As described before, sGAG content was determined

with a modification of the protocol described by Farndale et al.[47, 164]. Collagen content

was determined by an assay as described by Huszar et al. [165], and a standard curve was

prepared from trans-4-hydroxyproline (Sigma). The number of mature collagen

hydroxylysylpyridinoline (HP) and lysylpyridinoline (LP) crosslinks, as a measure of tissue

maturity were measured in the digests of the constructs using high performance liquid

Chapter 3

44

chromatography as described previously [189-191]. The number of HP and LP crosslinks

were expressed per triple helix (TH), and the ratio of HP and LP crosslinks was determined.


After 0 (only for unseeded group), 1, 2, 3, 4, 5, and 6 weeks of culture, the mechanical

properties of the engineered constructs (n=3-4 per group) were assessed by uniaxial

tensile tests in longitudinal direction of the constructs, using a BioTester 5000 (CellScale,

Canada). The samples were stretched to 5, 10, and 15% strain for 5 times to precondition

the samples. Mechanical test data was averaged per group. Sample thickness and width

were measured with an electronic caliper. The Young’s modulus was determined at a

strain of 15%.

3.3.6 Histology

To analyze tissue formation qualitatively, constructs were processed for histology (n=1

per group). Representative samples were fixed with 3.7% formaldehyde (Merck) and

embedded in paraffin. Tissue sections of 10 μm were cut and studied by Masson

Trichrome (MT) staining (MTC kit, Sigma, Venlo, Netherlands) for collagen deposition. The

stainings were analyzed using light microscopy (Axio Observer, Zeiss, Germany).


Statistics were performed using GRAPHPAD Prism (version 5.04) and differences were

considered significant for p-values <0.05. All data were presented as mean ± standard

error of the mean (SEM). Regression analyses were performed to determine changes in

tissue weight, scaffold weight, amount of ECM components, stiffness of the samples, and

crosslinks over time. In case of a significant in- or decrease, the percentual increase or

decrease was calculated using the predicted model equation. Also plateau and slope of

the different curves were compared using regression analyses. One-way ANOVA, followed

by a Tukey’s multiple comparison post-hoc test, was used to compare TE composition with

native tissues.

3.4 Results

3.4.1 Scaffold to tissue ratio changes over time

Dry weight of PCL scaffold material remained constant during culture time, while dry

weight of PGA-P4HB scaffolds indicated mass loss starting after week 1, with a decrease

of 93% compared to the initial values after 6 weeks of culture (p<0.05, Figure 3.1A). A

contribution in weight due to tissue formation was observed in both scaffold types, as

weight of tissues cultured in both PCL- and PGA-P4HB-based scaffolds increased during

culture. When comparing the ratio between tissue weight and remaining scaffold weight,


45

3

a percentual decrease in scaffold contribution and a percentual increase in tissue weight

was observed in both the PGA-P4HB-based (Figure 3.1B) and PCL-based (Figure 3.1C)

scaffold groups (p<0.05). After seeding, mainly scaffold weight was contributing to the

total weight of the constructs, as no tissue was formed yet. Although tissue was formed

within the PGA-P4HB constructs, a decrease in total weight was observed during culture,

which was due to the fast degradation, and thus mass loss, of PGA-P4HB scaffolds. In the

PGA-P4HB-based constructs a change was observed after roughly 2 weeks, as after this

time point mainly tissue weight contributed to the total weight of the constructs. PGA-

P4HB scaffolds were completely degraded after 6 weeks, with only tissue weight

contributing to the total weight of the constructs. This change was not observed in the

PCL-based constructs, as PCL scaffolds did not degrade as fast as PGA-P4HB scaffolds and

primarily contributed to the total weight.

Figure 3.1 Dry weights of scaffolds and tissues during culture (A). Weights of the scaffolds and tissues are given as mean ± SEM. PCL scaffold remains stable during culture, while PGA-P4HB started to degrade after 1 week already. Newly formed tissue is contributing to the total weight of the strips. Ratio of the fast-degrading PGA-P4HB (B) and the slow-degrading PCL (C) scaffold to tissue during culture, given in percentages. Weight of tissue and scaffold are given as mean percentage, as a section of the total weight of the samples. Total weight of the whole samples are set at 100%.

3.4.2 Tissue evolution in slow- and fast-degrading scaffold

Total amounts of DNA, sGAG and collagen per construct increased during culture in both

the PGA-P4HB (Figure 3.2A) and the PCL (Figure 3.2B) groups (all p<0.01). When

comparing the sGAG formation between the scaffold groups (Figure 3.2C), production

rates were similar, however the total amount formed of the PGA-P4HB-based tissues were

significantly lower compared to their PCL counterparts (p<0.01). Furthermore, although

Chapter 3

46

collagen production was increased in the PGA-P4HB groups compared to the PCL groups

during the first two weeks of culture, the total amount formed of the PGA-P4HB group

was lower compared to the PCL group after 6 weeks (p<0.01, Figure 3.2D). Total ECM

values in the PCL-based constructs after 6 weeks of culture were higher compared to the

PGA-P4HB-based constructs, which were 192±3 μg and 166±12 μg for the PCL and PGA-

P4HB constructs, respectively. Although lower total amounts were observed for the PGA-

P4HB-based tissues, cells in the PGA-P4HB-based tissues seemed to be more synthetic

during the first two weeks of culture compared to cells in PCL-based scaffolds, with

increased total amounts of ECM when corrected for the amount of DNA. Synthetic

activity, in terms of sGAG per DNA and collagen per DNA, was decreasing with time for

cells in the PGA-P4HB scaffolds (p<0.01), while this was increasing with time for the cells

in the PCL scaffolds (p<0.01) (Figure 3.2E).

Figure 3.2 Combined results of DNA, sGAG and collagen per strip during culture on PGA-P4HB (A) and PCL (B) scaffolds. During culture, the total amount of ECM increased, which was more pronounced for PCL-based tissues. PGA-P4HB-based constructs demonstrated lower plateau levels of amount of sGAG (C) and collagen (D) compared to PCL-based constructs. sGAG and collagen production per DNA (E) of PGA-P4HB-based tissues were increased during the first weeks, and became comparable to PCL-based tissues after 3 weeks. All results are given as mean ± SEM.


47

3

Both the HP and LP crosslinks per triple helix increased during cultured time in tissues

cultured on both the PCL and PGA-P4HB scaffolds (p<0.01, Figure 3.3A), with increased

production rates of HP per triple helix compared to LP per triple helix in both scaffold

groups (p<0.01). When compared between the scaffold groups, faster production of LP

per triple helix was observed in the PGA-P4HB-based tissues (p<0.05) compared to the

PCL-based tissues, while no difference in the production rates of HP per triple helix was

found between the scaffold groups. This resulted in a higher HP/LP ratio in the PCL-based

group (Figure 3.3B).

Figure 3.3 Collagen crosslinks in both scaffold groups given as HP/triple helix and LP/triple helix (A). Crosslinks within tissues grown on both type of scaffolds increased with culture time, while HP/triple helix increased with a higher rate compared to LP/triple helix. HP/LP ratio (B) was increased in PCL-based tissues compared to PGA-P4HB-based tissues.

3.4.3 Engineered tissues versus native heart valves

After 6 weeks of culture, the ECM amount per mg tissue (Figure 3.4A) and ratio (Figure

3.4B) of engineered tissues were compared to human aortic heart valves. DNA per mg

tissue was comparable between engineered tissue based on both scaffold types, and

native values of different age groups. sGAG per mg tissue is decreasing during ageing of

human, while the amount of collagen per mg tissue is increasing (p<0.05). All engineered

tissues resulted in sGAG values comparable to native adolescent and adult values, while

PGA-P4HB-based tissues demonstrated lower sGAG values compared to native values in

children (p<0.05). Collagen values of PCL-based tissues were not significant different from

native values, while PGA-P4HB-based tissues resulted in lower values (p<0.05 compared

to children, and p<0.001 compared to adolescents and adults). Although the amounts of

ECM differed between the engineered tissue groups, their ECM ratios were comparable.

These ratios were also similar to ratios found in children and adolescents. When

compared to adult tissues, the percentual portion of collagen differed between adults and

both engineered groups (p<0.05), while the percentage of sGAG was only significantly

different from the PCL-based tissues (p<0.05). Although the amount of newly formed

collagen in PCL-based engineered tissues after 6 weeks is similar to values measured in

native, the HP collagen crosslinks of both PCL and PGA-based tissues do not reach native

values during culture (data not shown). After 6 weeks of culture, HP crosslinks of PCL- and

PGA-P4HB-based tissues were 0.63±0.04 and 0.52±0.03 HP/triple helix, respectively, while

values observed in children, adolescents and adults were 2.0±0.1, 2.0±0.03 and 2.6±0.1

Chapter 3

48

HP/triple helix, respectively. LP/triple helix also showed to increase during culture time

and directed towards values measured in children (0.2±0.06 LP/triple helix) at the end of

culture. HP/LP ratio of engineered tissues was similar with native values for adolescents

and adults, however, differed significantly with the ratio found in children (p<0.001)

(Figure 3.4C). During aging, the ratio drops rapidly from 14.4 in children, towards 5.7 and

4.8 in adolescents and adults, respectively, as a result from a fast increase in LP/triple helix

in adolescents.

Figure 3.4 Comparison between amount of ECM (A) and ECM ratio (B) per mg formed engineered tissue with native data. Results are given as mean ± SEM. Tissues based on PCL scaffolds showed comparable amounts of ECM compared to native human aortic valve values, while amounts found in PGA-P4HB-based tissues were lower compared to their native counterparts. ECM ratio was similar in all engineered tissues, and comparable to ratios found in children and adolescents, while they differed compared to the ratio observed in adults. The HP/LP ratio (C) of the engineered tissues was comparable to the ratio observed in aortic valves of adolescents and adults, while HP/LP ratio was lower compared to children. #, * and ^ represent significant differences of sGAG, collagen and HP/LP ratio, respectively. Single or double symbols indicate p<0.05 and p<0.001.


49

3

3.4.4 Mechanical characteristics of formed tissues based on fast- or slow-degrading

scaffold

Due to fast loss of mechanical integrity of the PGA-P4HB scaffold strips, mechanical tests

on the unseeded PGA-P4HB samples over time could not be performed. This indicates that

the observed mechanical properties in the seeded PGA-P4HB constructs are solely

determined by the tissue only. Contribution of tissue formation to the mechanical

properties was observed in both seeded PGA-P4HB and PCL samples, as samples became

stiffer with culture time (p<0.01 for PCL and p<0.05 for PGA-P4HB, Figure 3.5), while the

Young’s modulus remained constant in the unseeded PCL scaffold strips.

Figure 3.5 Young’s modulus of seeded and unseeded scaffold strips during culture, given as mean ± SEM. Scaffold in PCL-based constructs is still contributing to the mechanical properties, while for the PGA-P4HB-based constructs mechanical properties are determined by tissue only. Newly formed tissue showed an additional effect on the Young’s-modulus, as demonstrated with an increased stiffness in the seeded samples compared to the unseeded samples.

3.4.5 Histological visualization of engineered tissues in time

Histology of the constructs revealed cellular tissues with dense surface layers, which was

more pronounced in PGA-P4HB-based tissues. Masson Trichrome stainings showed

collagen fibers throughout the strips of all groups during culture. Collagen is less

homogeneously distributed in the PGA-P4HB-based strips (Figure 3.6A-F) compared to the

PCL-based strips (Figure 3.6H-M). Furthermore, PCL-based tissues resulted in interstitial

growth of tissue, while appositional growth was observed in the PGA-P4HB-based tissues,

where a thick layer of tissue was formed around the scaffold. In addition, PGA-P4HB

constructs showed compaction (decreased scaffold width) during culture, with significant

differences compared to the original width (p<0.01), while the width of PCL constructs

remained stable during culture (Figure 3.6G).

Culture time [weeks]

E-m

od

ulu

s a

t 15%

str

ain

[M

Pa]

0 2 4 60

5

10

15PCL unseeded

PCL seeded

PGA-P4HB seeded

Chapter 3

50

Figure 3.6 Masson Trichrome staining of PGA-P4HB (week 1-6 representing by A-F) and PCL (week 1-6 representing by H-M) sections. The black scale bars represent 600 μm. Collagen is shown in blue, and red represents cytoplasm and muscle tissue. Vacuoles within the PCL sections are due to scaffolds remnants, which are dissolved during the dehydration step. PGA-P4HB sections do still show scaffold remnants (uncolored parts). Collagen is more homogeneously distributed in the PCL strips compared to the PGA-P4HB strips. Thickness of the strips (G) remains stable for PCL strips, while PGA-P4HB strips showed compaction.

3.5 Discussion

A balance between the rate of scaffold degradation and tissue formation is crucial for

maintaining mechanical integrity of the replaced tissues. We estimated the influence of

slow- versus fast-degrading scaffolds on the amount and composition of engineered

cardiovascular tissues, and mechanical integrity during culture. In addition, we compared

these values of the engineered tissues to values found in native human heart valves

leaflets.

3.5.1 In vitro evolution of tissue formation

The unseeded PCL scaffold strips did not degrade in vitro, in terms of weight, while the

unseeded PGA-P4HB scaffold started to loose mass already after 1 week. This resulted,


51

3

together with the contribution in weight of the tissues, in different scaffold to tissue ratios

during culture, with tissue weight being the main contributing factor in PGA-P4HB

constructs, while for PCL constructs both tissue and scaffold weight contributed to the

total weight. Our results on scaffold degradation are comparable with findings by Klouda

[153], where a mass loss of 0.9% and 11% for PCL and PGA-P4HB scaffolds, respectively,

was found after 15 days of static incubation. However, we observed a more severe mass

loss of PGA-P4HB scaffold, as it decreased by 33% after 14 days. This might be due to the

fact that in the study of Klouda samples were incubated with PBS, while our samples were

incubated in culture medium containing FBS. As certain enzymes present in serum are

known for degrading scaffolds [120, 123, 129, 192-194], this might have led to accelerated

degradation of the PGA-P4HB scaffold strips as compared to the study of Klouda.

Although increased amounts of ECM components were shown in both scaffold groups

with time, differences in tissue composition, when cultured using fast- or slow-degrading

scaffolds were observed. A first observation was that cells in the PGA-P4HB-based tissues

seemed to be highly synthetic as sGAG and collagen per DNA were higher compared to

PCL-based tissues. However, this was only observed during the first two weeks of culture,

where after the cells became less synthetic. At the end of culture, higher total amounts of

sGAG and collagen for PCL-based tissues were observed. We hypothesize that this

difference in tissue evolution is due to the fast degradation of the PGA-P4HB scaffolds,

resulting in highly synthetic cells during the first weeks to compensate for the loss of

mechanical integrity by newly formed collagen fibers. Compaction of the PGA-P4HB

scaffolds resulted in a smaller surface area and less volume for the cells within these

tissues to lay down their ECM, compared to PCL-based tissues, which might have resulted

in a higher amount of ECM after 6 weeks of culture in the latter.

It is well described that degradation of PGA scaffolds with or without P4HB can alter the

pH of the environment, due to their acid degradation products [181, 195-197]. A low pH

possibly affects the viability, proliferation or tissue synthesis of the cells. Higgins et al.

showed that the amount of porcine smooth muscle cells decreased and cells

dedifferentiated, due to PGA degradation products [195]. However, in their study media

was collected after 7 days only, while within our study media was changed twice a week

to prevent building up of degradation products and thus an acidic environment. We,

therefore, assume that released degradation products into the culture medium did not

have a profound effect on the viability, proliferation, and tissue synthesis within our

experiments.

All tissues demonstrated a continuous increase in LP and HP crosslinks during culture.

However, the ratio between these crosslink types differed between tissues. The HP/LP

ratio was lower for PGA-P4HB-based tissues compared to PCL-based tissues. Wassen et

al.[198] described that a lower HP/LP ratio caused by a relative high amount of LP/triple

helix, as observed in our PGA-P4HB samples, is seen in mineralized tissues only. This might

Chapter 3

52

assume that tissues cultured in PGA-P4HB scaffolds are more prone to mineralization

compared to their PCL counterparts. They also hypothesize that mineralization of collagen

fibrils is promoted by specific orientation of the molecules within these fibrils, which

might be different between the PCL and PGA-P4HB-based tissues due to a potential higher

degree of tissue remodeling of the PGA-P4HB-based tissues as a result of faster

degradation of PGA-P4HB scaffolds.

3.5.2 Comparison between engineered and native tissues

In literature, different methods are described to compare ECM components and amounts

of tissue, between engineered tissues, or to their native counterparts. These include a

non-invasive monitoring system to correlate biomarkers present in culture medium with

the synthesized tissue [187] and a method where ECM components are expressed as mg

per cm3 tissue [80]. However, these are suboptimal methods as the first one does not

include the total amount of tissue formed, and, in the second method, remaining scaffold

can contribute to the dimensions and, therefore, possibly influences the outcome,

especially when the scaffold is not degraded yet. To allow for accurate insight into tissue

evolution during culture and a fair comparison between engineered and native tissues,

only tissue weight without the contribution of remaining scaffold should be used. Our

study is the first that provides these insights as we corrected for the presence of remaining

scaffold. This correction is of importance when comparing properties of tissues that were

cultured using scaffolds with different degradation rates and when comparing engineered

tissues that were grown on slow-degrading scaffolds, which is still (partly) present, with

native tissues. A limitation of this method is that we do not account for the effect of cells

and tissue on scaffold degradation. The presence of cells can result in accelerated

degradation, as cells might release enzymes that stimulate this degradation. Furthermore,

in vivo macrophages will migrate to the scaffold materials and start to degrade the

materials, which is not the case in our in vitro set-up. Despite this limitation, this new

method comes nearest to the actual values, compared to all other studies performed until

today.

Figure 3.7 provides an overview of the generated amount of ECM and mechanical

properties, of PCL and PGA-P4HB-based tissues after 6 weeks of culture, compared to

values of pulmonary valves of children, which is the first target of tissue engineered

valves. PCL-based tissues show ECM values which are most similar to native values found

in children, while PGA-P4HB-based tissues showed a somewhat lower amount of ECM.

Similar stiffness values were observed in both PCL and PGA-P4HB-based tissues, while the

values of the latter is only determined by the newly formed tissue and not by remaining

scaffold, as observed for PCL. Stiffness of engineered samples are higher compared to

native values, however, we do not expect difficulties in opening or closing of the leaflets

after implantation, as PGA-P4HB valves with a similar stiffness were successfully

implanted before [68].


53

3

Figure 3.7 Comparison of native values of a child’s pulmonary valves with engineered tissues after 6 weeks of both PCL and PGA-P4HB-based tissues. Values of pulmonary valves of children are set at 100% (horizontal dashed line). Values of PCL and PGA-P4HB-based tissues are given as percentage compared to native values of children. PCL shows values that are similar or towards to native values in terms of ECM, while PGA-P4HB shows lower values. Engineered tissues are stiffer compared to their native counterparts. Although PGA-P4HB scaffold does not influence the mechanical properties of the tissue after 6 weeks, stiffness is similar to PCL-based tissues which are still partly influenced by remaining scaffold (marked area in the bar).

3.6 Conclusion

In conclusion, tissues based on slow-degrading materials, which maintained weight and

mechanical integrity during culture, demonstrated tissues which preserved their 3D

shape. Tissues based on fast-degrading material, which quickly demonstrated mass loss

and loss of mechanical integrity, resulted in compaction during culture and different tissue

to scaffold ratios. Although cells in PGA-P4HB constructs produced tissue at a higher rate

during the first weeks of culture compared to cells in PCL constructs, the amount of tissue

after 6 weeks was higher in the latter. ECM ratios were comparable between the scaffold

groups and also between engineered and native human values. This study demonstrates

the importance of using slow-degrading scaffolds in order to create constructs with stable

mechanical integrity, which maintain their configuration upon implantation. Further long-

term research is needed to investigate properties of PCL-based tissues when this scaffold

material is completely degraded.

Perc

en

tag

e

DNA

GAG

Colla

gen

Young's

-modulu

sDNA

GAG

Colla

gen

Young's

-modulu

s

0

50

100

150

200

250

PGA-P4HB

PCL

Pulmonary child

Chapter 3

54

Acknowledgements


Institute for Regenerative Medicine (NIRM, grant No. FES0908). This research also forms

part of the Project P1.01 iValve of the research program of the BioMedical Materials

institute, co-funded by the Dutch Ministry of Economic Affairs. The financial contribution

of the Nederlandse Hartstichting is gratefully acknowledged. The authors gratefully thank

Marc Simonet for electrospinning of the PCL scaffolds.

55

Hydrolytic and oxidative

degradation of electrospun

supramolecular biomaterials:

In vitro degradation pathways

4

M. Brugmans

S. Sontjens

M. Cox

A. Nandakumar

A. Bosman

T. Mes

H. Janssen

C. Bouten

F. Baaijens

A. Driessen-Mol

Submitted

Chapter 4

56

4.1 Abstract

The emerging field of in situ tissue engineering of load bearing tissues places high

demands on the scaffolds, as these scaffolds should provide mechanical stability

immediately upon implantation. A new class of synthetic biomaterials are the

supramolecular polymers, which contain non-covalent interactions between the polymer

chains, and can form complex 3D structures by self assembly. Here, we aimed to map the

degradation characteristics of promising (supramolecular) materials, as well as their

susceptibility to degradation. The selected biomaterials were all PCL, either unmodified

or with supramolecular (either 2-ureido-[1H]-pyrimidin-4-one or bis-urea units) hydrogen

bonding moieties incorporated into the backbone. As these materials contain elastomeric

properties, they are suitable for cardiovascular applications. Electrospun scaffold strips of

these materials were incubated with solutions containing enzymes that catalyze

hydrolysis, or solutions containing oxidative species. At several time points, chemical,

morphological, and mechanical properties were investigated. It was demonstrated that

conventional and supramolecular PCL-based polymers respond differently to enzyme-

accelerated hydrolytic or oxidative degradation, depending on the morphological and

chemical composition of the material. Conventional PCL is more prone to hydrolytic

enzymatic degradation as compared to the supramolecular materials, while the opposite

was shown when degraded by an oxidative pathway. Given this knowledge regarding

degradation characteristics of different (supramolecular) materials, we are able to tailor

degradation characteristics by combining different PCL backbones with additional

supramolecular moieties. This toolbox can be employed to screen, limit, and select

biomaterials for pre-clinical in vivo studies targeted to different clinical applications.


57

4

4.2 Introduction

Tissue engineering aims to restore tissue structure and function of diseased or damaged

tissues by implantation of specifically designed bioresorbable materials, with or without

the addition of cells [199-201]. Conventional tissue engineering aims to collect autologous

cells from patients, which are utilized for the in vitro generation of new tissues, and are

often cultured in bioreactors for several weeks before implantation. A new and promising

approach is in situ tissue engineering, in which in vitro culture is omitted and the patient’s

body is used as a bioreactor [86, 202-204]. New tissue will be regenerated directly in the

body by host cells after implantation of, for example, a bioresorbable electrospun

polymeric scaffold. This makes the overall procedure less demanding in terms of costs,

time, and regulatory challenges, and creates off-the-shelf availability.

In situ tissue engineering of load-bearing tissues places high demands on the

bioresorbable scaffolds, as these scaffolds should be able to provide mechanical stability

immediately upon implantation, and for a prolonged period thereafter, until sufficient

mature neo-tissue is formed by recruited cells to take over the mechanical function of the

scaffold. Various synthetic bioresorbable polymers are used for tissue engineering

applications, and these polymers include aliphatic polyesters (e.g. polylactic acid (PLA),

PGA and PCL), as well as various polyurethanes [129, 181, 205, 206]. A new set of synthetic

materials are the supramolecular polymers, which are formed by arrays of directed, non-

covalent interactions between the building blocks, and can form complex 3D-structures by

self assembly [112]. Material properties such as mechanics and resorption rate, which are

critical for the success of in situ tissue engineering can be modified by combining or

changing ratios of the same building blocks. This potentially allows for a variety of

polymers with varying properties to be synthesized in a relatively short time span, thereby

accelerating the development process. Monomeric units of the supramolecular polymers

possess a relatively low molecular weight, resulting in beneficial processing properties,

e.g. easy dissolution in organic solvents. Furthermore, supramolecular polymers may show

self-healing properties [113, 207, 208], can easily be made bioactive [119, 209], and allow

for a more controlled way of synthesis, which can result in complex molecular structures

[112]. Because of these features, these materials pose excellent candidates for use in in

situ tissue engineering. Particularly, we are interested in biomaterials that either have 2-

ureido-[1H]-pyrimidin-4-one (UPy) [114-117] or bis-urea (BU) [119] motifs incorporated

into their molecular structure, as these contain elastomeric properties, which makes them

suitable for cardiovascular applications.

To enable the formation of a completely autologous tissue, the scaffold should degrade at

the right pace during neo-tissue formation, leaving behind a living implant that is able to

remodel and grow. In vivo, degradation of implanted scaffold materials can be

accomplished via different pathways that operate at the same time, and that even may

affect each other [120, 121, 124, 128]. A well-known pathway is hydrolytic degradation,

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where chemical bonds of the polymer chains are cleaved by reaction with water

molecules, forming oligomers and ultimately generating small molecules that can be

cleared from the body [120, 121]. Previous studies have reported that several enzymes,

like proteases and esterases, which are present in human serum or are expressed by

macrophages and other activated cells that are in contact with the scaffold, are known to

catalytically accelerate this process [120, 123-125, 193]. Another well-described pathway

is oxidative degradation, which is mediated by ROS that are secreted by macrophages,

neutrophils and giant cells that are in contact with the scaffold [120, 126]. Previous studies

have investigated that oxidation of polymers is often initiated by abstraction of a

hydrogen atom by radicals, resulting in chain scission and/or crosslinking of the polymer

[210, 211]. Mapping the degradation characteristics of promising (supramolecular)

materials for use in in situ tissue engineering approaches, as well as their susceptibility for

certain degradation pathways, paves the way for screening and selection of materials for

various clinical implantation sites.

The degradation properties of widely used and well-known materials such as polyesters,

polycarbonates and polyurethanes have been examined extensively, both in vitro and in

vivo [125, 180, 206, 212-216]. In general, results of these studies show that polymers

containing ester or anhydride linkages react with water molecules and undergo hydrolysis

[58, 81, 121, 180]. The water molecules can access those chemical species more easily,

and thus increase the hydrolytic activity, when the polymer is amorphous or contains

aliphatic structures [121, 217]. Other polymers, including polyethers and polyurethanes,

were found to be more susceptible to the oxidative pathway, as these materials contain α-

methylene groups adjacent to ether or urethane groups, which are more prone to the

formation of carbon centered radicals by abstraction of a hydrogen atom [121, 127, 128,

131, 211, 218]. Just a few studies reported on the degradation characteristics of various

polymers (PCL, polycarbonates, or polyurethane) modified with UPy or BU units. These

were performed by incubating the materials in phosphate buffer saline (PBS) or solutions

of various lipases at 37°C [117, 209, 219, 220]. These studies showed that the rates of

enzymatic degradation can span a wide range, from less than 1% degradation after 1

month [220] to 90% after only 15 days [209], depending on the types of lipase and

polymers used. No hydrolytic degradation, in terms of weight loss, of the UPy containing

materials was observed for 126 days when samples were incubated with PBS [117], and a

decrease in weight of only 2% after 120 days was observed for BU-containing materials

[219].

Although these studies gave some insight into the degradation properties of bioresorbable

materials, the major part of these studies were performed on films or disks which are quite

dense, while degradation rate of electrospun scaffolds, that are more porous and have

higher surface to volume ratio, can be different. Studying the degradation properties of

electrospun meshes is, from a clinical point of view, more relevant as these are more likely

to be implanted as a tissue replacement, rather than a compact, solid construct.


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Furthermore, most research has focused on a single degradation pathway, while it is of

importance to assess either the enzyme-accelerated hydrolytic and the oxidative

degradation pathways, since in vivo both pathways may be operative and consequently,

both may affect the implanted scaffold.

Here, an in vitro study was designed to investigate both degradation pathways in an

accelerated fashion and was used to assess the degradation of several promising

supramolecular biomaterials for in situ tissue engineering. We have chosen three

previously reported supramolecular biomaterials, in which PCL backbones are combined

with either UPy hydrogen bonding groups (materials PCL2000-UPy and PCL800-UPy) [221] or

BU hydrogen bonding groups (PCL2000-BU) [119]. High molecular weight PCL, a material

frequently used for tissue engineering scaffolds, was added as a benchmark. All materials

were electrospun and the resulting scaffold meshes were either exposed to enzymes that

catalyze hydrolysis or to oxidative conditions. Degradation was monitored over time by

examining the remaining scaffold with respect to weight, molecular weight, fiber diameter,

and mechanical properties. Statistical analyses were performed to analyze changes in

properties over time of all polymers with the various treatments, as well as to investigate

their susceptibility to degradation and its mechanism (surface or bulk erosion).


4.3.1 Materials

All reagents, chemicals, materials, and organic solvents were obtained from commercial

sources and were used without further purification, unless otherwise noted. The PCL-

based supramolecular biomaterials PCL2000-UPy, PCL800-UPy and PCL2000-BU were

synthesized as previously described from PCL diol building blocks of molecular weights

800 or 2000 [119, 221]. These PCL2000-diol and PCL800-diol building blocks are prepared by

initiation from diethylene glycol, so they contain one ether bond in their structure.

Conventional PCL (Purasorb PC 12, IV=1.24 dl/g) was purchased at Purac Biochem,

Gorinchem, the Netherlands. Thermal characterization of these materials was performed

by differential scanning calorimetry (DSC) on a Perkin Elmer Pyris 1 or on a TA Instruments

Q2000. Reported data are from the melt, so after the sample has been in the isotropic

state, and were determined in the second heating run at a heating rate of 10°C/min. The

glass transition temperature (Tg) is reported as the inflection point, while the melting

transition (Tm) is reported as the peak of the transition.

4.3.2 Scaffold preparation

Scaffolds were fabricated in a climate-controlled electrospinning cabinet (IME

Technologies, Geldrop, The Netherlands) using the conventional electrospinning method

as described before [222]. Rectangular strips (25 (l) x 5 (w) x 0.44 (t) mm) were punched

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out of the electrospun scaffold meshes. Initial weight (W0) and thickness of all individual

strips were measured using a digital balance (Mettler Toledo, XS105, Greifensee,

Switzerland)) and a digital thickness gauge (Mitutoyo, SGM, Groningen, The Netherlands).

Prior to incubation for degradation, the meshes were centrifuged at 4500 rpm in purified

water for 5 minutes to remove air bubbles.

4.3.3 Accelerated in vitro degradation

Strips (n=60 per material) were incubated at 37°C in 1.5 ml enzyme solution, referred to

as enzymatic degradation, or in a 4 ml oxidative degradation solution each. The enzyme

solution consisted of 100 U/mL lipase from Thermomyces lanuginosus (L0777, Sigma-

Aldrich) in PBS or 10 U/mL cholesterol esterase from bovine pancreas (C-3766, Sigma-

Aldrich) in PBS. These enzymes, which are present in serum and are secreted by activated

macrophages, are known to cleave ester and urethane bonds to a higher extent as

compared to other secreted enzymes [215, 223, 224]. The oxidative solution comprised

of 20% hydrogen peroxide (Sigma-Aldrich) and 0.1 M cobalt(II) chloride (Sigma-Aldrich) in

purified water (pH of this solution is 4.5). Hydrogen peroxide and cobalt(II) chloride

undergo a Haber-Weiss reaction, creating reactive hydroxyl radicals [211]. Incubation

times of the scaffolds in lipase, cholesterol esterase, or oxidative solutions were up to 56,

96 and 400 hours, respectively. Based on literature [131, 225], solutions were changed

every 3-4 days to maintain enzymatic activity and a constant concentration of radicals.

4.3.4 Scaffold characterization

Analyses of the (remaining) scaffolds were performed at 5 time points for the enzymatic

groups and 7 time points for the oxidative group (n=4 per group per time point). Mass

loss, molecular weight, fiber diameter, and mechanical properties were determined.

4.3.4.1 Mass loss

Scaffold strips were removed from the degradation solution, washed three times with

purified water, dried under vacuum at 37°C for 16 hours and weighed (Mettler Toledo,

XS105, Greifensee, Switzerland), to assess weight loss due to scaffold degradation. Mass

loss of the scaffolds (n=4 per group per time point) was determined using the equation:

W1/W0 × 100%, where W0 is the initial scaffold weight and W1 indicates remaining scaffold

weight.

4.3.4.2 Scanning electron microscopy (SEM)

Scaffold fiber morphology and average fiber diameters were assessed and determined by

scanning electron microscope (SEM), (Phenomworld, Eindhoven, The Netherlands) of one

sample per group per time point. Average fiber diameters were determined by 20


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individual measurements performed on four SEM images per scaffold strip, using

Phenomworld software (Fibermetric, Phenom pro suite version 2.0).

4.3.4.3 Mechanical properties

To study the effect of degradation on the mechanical properties of the scaffolds, uniaxial

tensile tests in longitudinal direction of the strips (n=3 per group per time point) were

performed. Due to a loss of mechanical integrity over time, associated with degradation,

it was not possible to perform tensile tests on all PCL-BU and PCL-UPy strips of the latest

oxidation time points. Sample thickness and width were measured with an electronic

caliper. Stress-strain curves were obtained (Mecmesin multiTest-i) at an elongation rate

of 100% per minute and the mechanical test data was averaged per group per time point.

The elasticity modulus (Young’s-modulus) was determined as the slope of the initial linear

part of the curve, as a measure for stiffness. As a measure for strength, the ultimate

tensile strength (UTS) was defined as the peak stress value, while strain at break is a

measure for the maximal elongation of the samples until break.

4.3.4.4 Molecular weight (GPC)

After tensile testing, one strip per group per time point of each material was taken and

dissolved in dimethylformamide ((DMF), Sigma) in order to determine the mass averaged

molecular weight (Mw) of the samples by gel permeation chromatography (GPC) analysis.

GPC was performed on a Varian/Polymer Laboratories PL-GPC 50, using DMF with 10

mmol/L lithium bromide as eluent and maintaining the temperature of the equipment at

50°C. The relative or apparent molecular weights (Mw) were determined with respect to

polyethylene glycol standards. Samples were measured in duplicate and the Mw was

averaged from this duplicate measurement.


All data are presented as mean ± standard deviation. Statistics were performed using

GRAPHPAD Prism (version 5) and differences were considered significant for p-values

<0.05.

4.3.5.1 Changes over time

Regression analyses were performed to determine changes in weight, Mw, fiber diameter,

Young’s-modulus, UTS, and strain at break over time. Both a one-phase decay model

(assuming the rate at which changes occur is proportional to the amount that is left) and

a linear model (assuming a constant rate) were used to fit the data. In case of a significant

increase or decrease with p<0.05 or p<0.01, the percentual in- or decrease was calculated

from the predicted model equation and classified as non-relevant (0-10%), small (10-

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25%), moderate (25-100% for an increase and 25-50% for a decrease), and severe (>100%

for an increase and 50-100% for a decrease).

4.3.5.2 Susceptibility to degradation and its mechanism

The susceptibility for both enzymatic and oxidative degradation was determined via

correlation analyses of all measured parameters. Significant correlations were classified

as weak (p<0.05), average (p<0.01), and strong (p<0.001). Susceptibility for degradation

was calculated as the number of significant correlations (with more weight to the average

and strong correlations as compared to the weak correlations) divided by the maximum

number of possible correlations and expressed as a percentage. Susceptibility was

classified as not susceptible (<20%), susceptible (20-60%), or highly susceptible (>60%).

To obtain insight into the mechanism of degradation, correlations were either attributed

to surface erosion or to bulk erosion. Correlations that were considered to attribute to

surface erosion were correlations between mass loss and fiber diameter, between

mechanical properties, between mechanical properties and fiber diameter, and between

mass loss and mechanical properties. Correlations that were considered to attribute to

bulk erosion were correlations between Mw and mass loss, mechanical properties, or

fiber diameter and inverse correlations between parameters. The susceptibility to either

enzymatic or oxidative degradation was subsequently determined as described above

with similar classifications for susceptibility.

4.4 Results

4.4.1 Material properties

The studied supramolecular biomaterials PCL2000-UPy, PCL800-UPy and PCL2000-BU are in

fact thermoplastic elastomers with PCL soft blocks and hard blocks composed of

interacting and phase separated hydrogen bonding units (Figure 4.1).

Figure 4.1 Schematic overview of the materials examined in this degradation study, with bis-urea (BU) (A), and ureidopyrimidinone (UPy) (B) based supramolecular biomaterials.


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PCL is a semi-crystalline polyester (Tg = –64°C, Tm = 52°C), while the PCL2000-BU

thermoplastic elastomer shows a first melting transition (Tm1) of the semi-crystalline PCL

soft block at a lower temperature and a second melting transition (Tm2) of the BU hard

block at a higher temperature (Tg = –54°C, Tm1 = 27°C, Tm2 = 98°C) [119]. Both PCL800 and

PCL2000–UPy are also thermoplastic elastomers (PCL800-UPy: Tg = -39°C, Tm1 = 65°C, Tm2 =

91 °C; PCL2000-UPy: Tg = -58°C, Tm1 = 53°C, Tm2 = 116°C).

4.4.2 In vitro degradation as monitored by scaffold mass loss and Mw

Enzymatic degradation (Figures 4.2A-D) of conventional PCL scaffolds resulted in

moderate (44%, p<0.01) to severe (92%, p<0.01) mass loss by lipase and cholesterol

esterase treatment, respectively, while Mw remained constant over time. For the

supramolecular materials, only the PCL2000-BU was affected by enzymatic degradation

with moderate weight loss by both lipase (30%, p<0.01) and cholesterol esterase (22%,

p<0.01) treatment (Figures 2A,C). Mw of PCL2000-BU did not change with lipase treatment,

while a small decrease in Mw (14%, p<0.05) was observed during cholesterol esterase

treatment. The PCL-UPy materials did not show changes in weight and Mw over time due

to enzymatic degradation.

Oxidative degradation (Figures 4.2E,F) did not affect mass and Mw of conventional PCL

scaffolds, while all supramolecular materials were affected. Both PCL-UPy materials

showed moderate mass loss (42% and 27%, p<0.01 for PCL800-UPy and PCL2000-UPy,

respectively) and a severe reduction in Mw (71% and 83%, p<0.01 for PCL800-UPy and

PCL2000-UPy, respectively). The PCL2000-BU also demonstrated moderate mass loss (35%,

p<0.01) and a severe reduction in Mw (94%, p<0.01) due to oxidative degradation.

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Figure 4.2 Influence of enzymatic (A-D) and oxidative degradation (E,F) on mass loss (A,C,E) and Mw (B,D,F) of conventional and supramolecular PCL-based scaffold strips. Significant and relevant changes over time are indicated by lines between data points. Conventional PCL was mainly affected by enzymatic degradation with moderate to severe mass loss, but with stable molecular weight. The supramolecular materials were mostly affected by oxidative degradation, with mass loss as well as decreases in molecular weight.

4.4.3 In vitro degradation as monitored by scaffold fiber diameter and morphology

Enzymatic degradation (Figures 4.3A,B) of conventional PCL scaffolds resulted in small to

severe fiber diameter reduction, depending on the enzyme used (18% and 62%, p<0.01

for cholesterol esterase and lipase treatment, respectively). Enzymatic degradation did not

affect the fiber diameter of both PCL-UPy materials, but resulted in a moderate reduction

in fiber diameter in PCL2000-BU scaffolds after both lipase (31%, p<0.05) and cholesterol

esterase (25%, p<0.01) treatment.


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Oxidative degradation (Figures 4.3C) did not affect the fiber diameter of the conventional

PCL scaffolds, but resulted in moderate reduction in fiber diameter for the PCL-UPy

materials (45% and 49%, p<0.01 for PCL800-UPy and PCL2000-UPy, respectively). PCL2000-BU

showed a small reduction in fiber diameter after oxidative treatment (10%, p<0.01).

Figure 4.3 Influence of enzymatic (A,B) and oxidative degradation (C) on the fiber diameter of conventional and supramolecular PCL-based scaffold strips. Significant and relevant changes over time are indicated by lines between data points. The fiber diameter of conventional PCL scaffolds was affected only by enzymatic degradation, while the supramolecular materials showed mainly reduced fiber diameters with oxidative degradation.

SEM images of scaffold strips before and after enzymatic and oxidative degradation

treatment confirmed these changes in fiber diameter (Figure 4.4). They further

demonstrate that the surface of the conventional PCL fibers is clearly affected by

degradation, while the fiber surface of the supramolecular materials seemed less affected

as compared to the conventional PCL, though more fragmented fibers were observed in

the supramolecular scaffold groups.

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Figure 4.4 SEM images with different magnifications of PCL-based scaffold strips before (A-D) and after enzymatic (E-L) and oxidative (M-P) degradation. Conventional PCL is mainly affected by enzymatic degradation, resulting in thinner and clearly affected fibers, while the supramolecular materials were mainly affected by oxidative degradation with thinner fibers. The fiber surface of the supramolecular materials seemed less affected as compared to the conventional PCL, though more fragmented fibers were observed. The white dots on the conventional PCL scaffold after oxidative degradation are presumably cobalt chloride remnants. White scale bars represent 20 micrometer.

4.4.4 Changes in mechanical properties during in vitro degradation

Enzymatic degradation (Figure 4.5A, C, E) resulted in overall weakening of conventional

PCL scaffolds with severe reductions in Young’s modulus (96%, p<0.01 and 57%, p<0.05

for lipase and cholesterol esterase, respectively), UTS (96%, p<0.05 and 51%, p<0.01 for

lipase and cholesterol esterase, respectively), and strain at break (80% and 66%, p<0.05

for lipase and cholesterol esterase, respectively). The PCL-UPy materials did not

demonstrate weakening, but changed mechanical properties with small to moderate

increases in modulus, depending on the PCL soft segment length, for both lipase (13% and

44%, p<0.01 for PCL2000-UPy and PCL800-UPy, respectively) and cholesterol esterase

treatment (19% and 99%, p<0.05 for PCL2000-UPy and PCL800-UPy, respectively). PCL2000-

UPy further showed a moderate reduction in strain at break with lipase treatment (27%,

p<0.05), indicating a change towards a more brittle material. PCL2000-BU showed a

combination of weakening and a change toward a more brittle material with moderate


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reductions in UTS (40%, p<0.01) and strain at break (39%, p<0.01) by cholesterol esterase

treatment, and a severe increase in modulus (56%, p<0.01) after lipase treatment.

Oxidative treatment (Figures 4.5B, D, F) only affected strain at break of conventional PCL

scaffolds with a moderate decrease (25%, p<0.05), while modulus and UTS remained

unaffected. The PCL-UPy materials showed a combination of weakening and a change

toward a more brittle material with severe reductions in UTS (96%, p<0.05 and 87%,

p<0.01 for PCL800-UPy and PCL2000-UPy, respectively) and strain at break (100% and 96%,

p<0.01 for PCL800-UPy and PCL2000-UPy, respectively), and a severe increase in Young’s

modulus (>300%, p<0.01) after lipase treatment. Similar weakening and changes toward

a more brittle material were observed for PCL2000-BU with severe reductions in UTS (80%,

p<0.05) and strain at break (99%, p<0.05), accompanied by a severe increase in Young’s

modulus (>1000%, p<0.01) after oxidative treatment.

Figure 4.5 Influence of enzymatic (A,C,E) and oxidative (B,D,F) degradation on the Young’s modulus (A,B), UTS (C,D), and strain at break (E,F) of PCL-based scaffold strips. The results for cholesterol esterase treatment are not shown, but are comparable to those of the lipase treatment. Significant and relevant changes over time are indicated by lines between data points. The mechanical properties of conventional PCL were mainly affected by enzymatic degradation and represented by overall weakening. The mechanical properties of the supramolecular materials were affected by enzymatic degradation, but to a larger extent by oxidative degradation. Here, a change to a more brittle material was evident, accompanied by an overall weakening of the material.

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4.4.5 Susceptibility to degradation and its mechanisms

Correlation analyses revealed that the conventional PCL scaffolds were susceptible to

enzymatic degradation, with the degree of susceptibility depending on the enzyme used.

Susceptibility was higher for lipase (62%) as compared to cholesterol esterase (36%) and

surface erosion seemed the dominant degradation mechanism (77% and 33% for lipase

and cholesterol esterase treatment, respectively). Conventional PCL scaffolds were not

susceptible (13%) to oxidative degradation.

The PCL-UPy materials were not susceptible to enzymatic degradation, neither lipase (2%

for both PCL800-UPy and PCL2000-UPy) nor cholesterol esterase (7% for both PCL800-UPy and

PCL2000-UPy). The susceptibility for oxidative degradation was dependent on the PCL soft

segment length, with no susceptibility for PCL800-UPy (13%) to susceptible for PCL2000-UPy

(40%). Both surface erosion (23%) and bulk erosion (33%) seemed involved. PCL2000-BU

was susceptible to enzymatic degradation, though only to cholesterol esterase (16% and

31% for lipase and cholesterol esterase, respectively), and oxidative degradation (24%).

Surface erosion seemed the dominant mechanism in degradation of PCL2000-BU (33% and

23% for enzymatic and oxidative degradation, respectively).

4.5 Discussion

Electrospun bioresorbable scaffold meshes represent promising candidates for use in in

situ tissue engineering to replace diseased or damaged tissue parts. While providing initial

mechanical stability upon implantation, host cells are recruited over time for neo-tissue

formation, taking over the mechanical function of the scaffold. Supramolecular polymers

represent interesting candidates to replace soft and elastic dynamically-loaded tissues. To

allow for the development of a stable fully autologous implant that can grow and remodel

in the patient, the scaffold should degrade in pace with neo-tissue formation. Here, we

aimed to map the degradation characteristics of promising (supramolecular) materials, as

well as their susceptibility to certain degradation pathways, for use in in situ tissue

engineering approaches. An in vitro test was designed to investigate the degradation of

electrospun biomaterial scaffolds either by enzyme-accelerated hydrolysis or by

oxidation. In addition to changes in fiber morphology of the meshes, changes in mass of

the scaffold, and changes in molecular weight, this in vitro study also monitored and

assessed changes in the mechanical properties of electrospun scaffolds over time. The

investigated scaffolds were prepared from PCL-based supramolecular biomaterials and

conventional PCL served as a benchmark. Figure 4.6 provides a summary of the results

obtained in this study indicating the changes over time with both enzymatic and oxidative

degradation as well as their susceptibility for each polymer group.


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Figure 4.6 Schematic summary of the obtained results indicating the changes over time in mass, Mw, fiber diameter, and mechanical properties by either enzymatic or oxidative degradation over time for conventional and supramolecular PCL-based scaffolds. Further, the susceptibility of each polymer group to enzymatic as well as oxidative degradation is represented by a color scale, with red indicating a high susceptibility and green referring to the material being not susceptible to degradation. Illustrations by Anthal Smits

The conventional PCL scaffolds were rapidly degraded by enzymatic hydrolysis, using

lipase or cholesterol esterase, as evidenced by mass loss, changes in fiber morphology,

and overall weakening, while molecular weight remained unaffected. These results are

consistent with findings by others, although different types and concentrations of the

enzymes resulted in slower or faster degradation of the PCL [125, 193, 226-229]. Polymer

degradation by enzymes can be either surface erosion or bulk degradation, depending on

the accessibility of the interior of the polymer to the enzyme. Surface erosion was

identified here as the dominant degradation mechanism with clear effects to the fiber

surface, thus apparently, the ability of the enzymes to infiltrate the hydrophobic semi-

crystalline PCL is limited (or the activity of the enzyme becomes compromised upon

infiltration) [120, 130].

In contrast to the conventional PCL meshes, the supramolecular UPy- and BU-based PCL

demonstrated to be less prone to hydrolyze enzymatically with no or minimal changes in

mass, molecular weight and fiber diameter. The PCL-UPy materials were classified as not

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susceptible to enzymatic degradation, though they demonstrated an increased Young’s

modulus, accompanied by a reduction in strain at break, which indicates a change to a

more brittle material. This change may be caused by annealing of the material at 37°C,

resulting in a material with an increased crystallinity of the PCL phase, and thus a more

brittle material. The PCL-BU was classified as susceptible to cholesterol esterase and not

to lipase, though also with a change to a more brittle material. Clearly, the introduction

of the BU or UPy hard blocks in the PCL backbone has a marked stabilizing effect on the

enzymatic degradation rate, despite increasing the overall polarity of the biomaterial by

introduction of the polar BU or UPy groups. Presumably, the different morphology of the

materials as compared to conventional PCL is causing the changes in hydrolytic enzymatic

degradation behavior. For PCL-BU, it is known that phase separation of the PCL soft block

and the BU hard block is on the nanometer scale (ca. 10 nm scale) [119], implying that the

partly amorphous PCL soft block in PCL-BU may be less accessible as compared to the

more sizable amorphous PCL phases in conventional semi-crystalline PCL. Moreover, the

molecular dynamics of the segmented PCL chains may be compromised, first as these

chains are relatively short, and second as they are kept into position at both ends by the

immobile UPy or BU hard blocks. According to the above factors, we propose that the PCL

chains in the supramolecular biomaterials are less accessible to enzymes, and therefore

causing the lowered enzymatic degradation susceptibility. Among the supramolecular

materials, the PCL-BU was more susceptible to enzymatic degradation as compared to

PCL-UPy, though similar PCL soft segment length were used in the backbone of PCL2000-

UPy and PCL2000-BU. Apparently, the ester bonds in the BU-based material are more

accessible and/or prone to hydrolysis as compared to those in the UPy containing

material. Both materials have phase separated soft and hard blocks and the exact manner

in which this phase separation takes place may influence and determine their

degradation. However, the exact differences in morphology, e.g. level of phase

separation, mobility of the PCL chains, and the level of crystallinity of the PCL soft phase,

between PCL-BU and PCL-UPy are not known and should be further investigated.

Oxidative degradation gave the opposite result as that observed for the enzyme-

accelerated degradation. Conventional PCL scaffolds were not susceptible to oxidative

conditions, with stable mass, molecular weight, fiber diameter, Young’s modulus, UTS,

and only a small decreased strain at break. The presence of merely an aqueous solution

without enzymes was clearly not enough to hydrolyze conventional PCL. Conventional PCL

only has ester groups in its structure, and apparently these ester groups are not

significantly degraded by the offered oxidative cobalt (II) solution, despite the fact that

the amorphous phase in semi-crystalline PCL must be accessible to the presented small

oxidative cobalt (II) and H2O2 (derived) species. The supramolecular materials on the other

hand did show susceptibility to oxidative degradation, with decreases in mass, molecular

weight, fiber diameter and a combination of weakening of the materials with a change to

a more brittle material, with some fragmentation of fibers. We primarily attribute the


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augmented oxidative degradation of the supramolecular PCL-based biomaterials to their

chemical differences with PCL (and not to morphological features), whereby these

differences are represented by the presence of ureidopyrimidone for PCL-UPy and the

presence of urea groups for PCL-BU. Moreover, all supramolecular PCL are based on PCL-

diols initiated from diethylene glycol, hence they comprise a single ether group in every

PCL soft block, which also might result in an increased sensitivity to oxidative degradation.

Remarkably, the PCL soft segment length influenced the susceptibility to oxidative

degradation, with the PCL800 soft block providing more resistance to oxidative

degradation. Oxidative degradation of the supramolecular materials was classified as

surface erosion, though for the PCL-UPy bulk erosion was also noted, indicating diffusion

of the small oxidative cobalt and H2O2 (derived) species into these materials.

It has to be noted that degradation is a dynamic process, as the mechanical, morphological

and chemical properties of the polymers change during degradation, and all can affect

surface and bulk mobility, accessibility by enzymes, and the diffusion of small molecules

such as water, oxidative species and degradation products. Here, we have investigated

degradation by enzymatic hydrolysis and oxidation separately. Future studies should

include a combination of degradation pathways to assess their combined effects.

Macrophages play an essential role in the degradation of polymeric scaffold meshes in in

situ tissue engineering as an inflammatory response provides the basis for neo-tissue

formation. These macrophages secrete both enzymes as well as oxidative species,

therewith triggering both degradation pathways. We are currently investigating whether

macrophage phenotype, claimed essential in tissue outcome [230], is correlated to trigger

degradation in either of the pathways, enabling assessment of the desired degradation

characteristics of electrospun bioresorbable meshes.

Depending on the application, either fast or slow resorption by the body is desired. When

scaffold resorption is too slow it can result in stress shielding of the growing tissue,

thereby impeding the regeneration process [231] or leading to undesirable outcomes.

When the resorption process is too fast, the mechanical integrity of the implant is not

sufficient, as the neo-tissue is not sufficiently developed yet to bear the full mechanical

force required [232], leading to failure of the implant. Furthermore, the site of

implantation might influence the resorption rate of a biomaterial. Mechanical forces, like

compression, fatigue and shear stress, or external factors like pH might affect the

resorption rate of the implanted material [121]. This again demonstrates the need to

tailor the properties of bioresorbable polymers specifically to the intended application.

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4.6 Conclusion

In this study, we demonstrated that conventional and supramolecular PCL-based

polymers respond differently to in vitro enzyme-accelerated hydrolytic or oxidative

degradation pathways, depending on the morphological and chemical composition of the

material. Conventional PCL is more prone to hydrolytic enzymatic degradation as

compared to the supramolecular materials, while the opposite is shown for these

materials when degraded by an oxidative pathway. Given this knowledge on degradation

characteristics of different (supramolecular) materials, we are able to tailor degradation

characteristics by combining different PCL backbones with additional supramolecular

moieties. This toolbox can be employed to screen, limit and select biomaterials that are

going to be used for pre-clinical in vivo studies for different clinical applications.

Acknowledgements

This research forms part of the iValve project of the research program of the BioMedical

Materials institute, co-funded by the Dutch Ministry of Economic Affairs, Agriculture and

Innovation. The financial contribution of the Nederlandse Hartstichting is gratefully

acknowledged. Part of the work by Marieke Brugmans was supported by a grant from the

Dutch government to the Netherlands Institute for Regenerative Medicine (NIRM, grant

No. FES0908). The authors would like to thank Roel Lalieu and Nanayaa Bates from Xeltis

for producing the scaffold sheets. Furthermore, Leonie Grootzwagers and Anita van de

Loo from Xeltis are acknowledged for performing the tensile tests and taking the SEM

images.

73

Advanced electrospun scaffold

degradation by inflammatory

macrophages in comparison with

healing macrophages

5

M. Brugmans

M. Cox

C. Bouten

F. Baaijens

A. Driessen-Mol

In preparation

Chapter 5

74

5.1 Abstract

Implantation of a synthetic scaffold made from a bioresorbable material will cause an

inflammatory response in which macrophages are claimed to play an essential role.

Macrophages show a continuum of functional properties alternating between the pro-

inflammatory (M1) and the tissue-healing (M2) phenotypes. The contribution of

macrophage phenotype to biomaterial resorption remains unclear and should be further

elucidated to provide more insight into the immune response to implanted biomaterials,

which is of particular relevance for in-situ tissue engineering approaches. In this study, 2D

and 3D in-vitro cultures were used to investigate the contribution of macrophage

polarization to degradation of electrospun biodegradable scaffolds. In addition, we

monitored the phenotypical change of unpolarized macrophages during time as an

indication of the initial macrophage response to the electrospun meshes. Monocytes of

the human cell line THP-1 were differentiated towards macrophages, seeded into 6-wells

plates (2D) or onto rectangular electrospun PCL strips (3D), and polarized towards

inflammatory (by LPS/IFNɣ) or healing macrophages (by IL4/IL13), or kept unpolarized. In

2D cultures, sample groups were sacrificed after 1, 3, 6, 8, and 10 days and cells were

counted. Furthermore, cell phenotype was assessed from cell morphology via imaging

before sacrifice. The 3D samples were sacrificed after 2 days, 1 week, and 4 weeks.

Samples were assessed with respect to DNA content, microstructure (SEM, with and

without cells), esterase activity, and gene expression (qPCR). Different cell morphologies

were observed between the polarized groups, whereas DNA amount decreased with time

for all phenotypes in both 2D and 3D cultures, albeit more prominent in the LPS/IFNɣ

polarized cells. Unpolarized cells demonstrated similar gene expression levels compared

to the healing phenotype. Scaffold degradation was observed in all phenotype groups, but

was most pronounced by the LPS/IFNɣ polarized cells. These findings were confirmed by

esterase activity and gene expression analysis. In conclusion, macrophage phenotype

affects the rate of electrospun scaffold degradation, with inflammatory macrophages

accelerating degradation.

Contribution of macrophage phenotype to scaffold degradation

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5

5.2 Introduction

The use of scaffold materials composed of bioresorbable synthetic polymers is a promising

approach to replace diseased tissues in patients, as these biomaterials are supposed to

be resorbed while new tissue is formed simultaneously. Implantation of a biomaterial

evokes an inflammatory response, in which macrophages play an essential role. These

cells provide the basis for neo-tissue formation [230], as well as degradation and removal

of the implanted polymeric scaffolds [126, 126, 233]. Macrophages show a continuum of

functional properties alternating, dependent on micro-environmental factors, between

the pro-inflammatory (M1) and the tissue-healing (M2) phenotypes [136, 234]. M1

macrophages are driven by pro-inflammatory signals, such as interferon gamma (IFN-ɣ)

and lipopolysaccharide (LPS) and secrete pro-inflammatory cytokines and ROS, while M2

macrophages are driven by interleukins (IL), such as IL4/IL13, and are involved in wound

healing and anti-inflammatory processes in favor of ECM formation [139, 235].

Although several studies focused on the suppression of the inflammatory response to

improve the biocompatibility of an implanted material [236-239], this inflammatory

response, in particular the balance between the phenotype of macrophages, is believed

to play an important role in the final outcome of tissue regeneration (balance towards M2

phenotype) or chronic inflammation and scar formation (balance towards M1 phenotype)

[135, 136, 140, 235, 240]. In a functional healing process, this balance is desired to be

mainly a M2 phenotype during the regenerative phase of wound healing.

Single macrophages are able to phagocytose small foreign body particles (<10 um) [138],

while larger particles (10-100 um) are beyond the phagocytic capacity of a single

macrophage and are phagocytized by fused macrophages, which form multinucleated

foreign body giant cells (FBGC). In case materials with sizes in the millimeter range are

implanted even the FBGC are not capable to engulf these large bulk materials and,

therefore, undergo ‘frustrated phagocytosis’, whereby ROS and enzymes are released in

an attempt to degrade the scaffold material. In a previous study, we used in vitro

degradation models to demonstrate the effect of enzymes or ROS products on different

(supramolecular) PCL-based scaffold materials (Brugmans et.al, 2015, submitted). It was

demonstrated that both enzymes and/or ROS products are able to degrade scaffold

materials, depending on the type of biomaterial. Pro-inflammatory M1 macrophages are

known to secrete both enzymes and ROS products [126, 139, 233], suggesting that M1

macrophages play an important role in scaffold degradation in vivo. Different studies have

shown that the interleukins IL-4 and IL-13 result in formation of FBGC [126, 241], and are

also known to polarize macrophages to the anti-inflammatory M2 phenotype [242]. FBGC

formation by IL-4 and IL-13 suggests that M2 macrophages are also involved in scaffold

degradation. Furthermore, differences in gene expression between M1 and M2 polarized

macrophages were studied. Results showed that both the M1 and M2 phenotypes secrete

enzymes, such as lipase A cholesterol esterase, which are known to be able to degrade

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76

scaffold materials that contain ester bonds [223, 243]. Taken together, the contribution

of macrophage phenotype in scaffold degradation remains unclear and should be further

elucidated to provide more insight into the desired immune response to implanted

biomaterials toward functional healing. Here, we investigated if and how macrophage

phenotype, claimed essential in tissue outcome, contributed to the degradation of

electrospun bioresorbable meshes. For this purpose, we made use of in vitro models

which included cultures of macrophages, polarized towards the inflammatory or healing

phenotype on two-dimensional substrates, or on electrospun scaffolds (3D). In addition,

unpolarized macrophage were cultured to investigate the contribution of the scaffold

meshes on the phenotypical change of these macrophages, as an indication for the initial

response of macrophages to the scaffold meshes.



Conventional PCL (Purasorb PC 12, IV=1.24 dl/g) was purchased at Purac Biochem,

Gorinchem, the Netherlands. Scaffolds were fabricated in a climate-controlled

electrospinning cabinet (IME Technology, Eindhoven, The Netherlands) using the

conventional electrospinning method, as described before [222]. Initial fiber morphology

was determined by scanning electron microscopy (SEM) (Phenomworld, Eindhoven, The

Netherlands). Rectangular strips (12.5 (l) x 5 (w) x 0.30 (t) mm) were punched out of

electrospun scaffold meshes. Strips were sterilized by gamma irradiation before use

(Synergy health, Ede, The Netherlands).

5.3.2 Cell culture and seeding

5.3.2.1 Cell culture

The human monocytic cell line THP-1 was purchased from Cell Lines Service (CLS,

Eppelheim, Germany) and cultured according the suppliers’ recommendation. The cells

were cultured in Roswell Park Memorial Institute (RPMI) 1640 medium with L-Glutamine

and 25 mM HEPES (Invitrogen, Breda, The Netherlands), supplemented with 10% Fetal

Bovine Serum ((FBS), Greiner Bio one, Frickenhausen, Germany), 1%

Penicillin/Streptomycin (P/S, Lonza, Basel, Switzerland), and 0.05mM 2-mercaptoethanol

(Sigma Aldrich) in a humidified atmosphere containing 5% CO2 at 37°C. Cell densities were

maintained between 0.5-1.0*106 cells per ml medium. Medium was changed 3 times per

week.


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5

5.3.2.2 Seeding of monocytes and transformation into macrophages

The experimental design of our study is shown in Figure 5.1. Scaffold strips (n=77) were

placed in 50 ml tubes containing 10 ml sterile PBS (Fisher, Landsmeer, The Netherlands).

These tubes were centrifuged for 10 minutes at 3500 rpm to increase hydrophilicity of the

scaffolds. Upon seeding, the scaffolds (n=72) were placed into 2 ml vials containing 1.5 ml

culture medium, 3.0*106 THP-1 monocytes, and 50 ng/ml of phorbol 12-myristate 13-

acetate ((PMA), Sigma Aldrich) to transform the monocytes into macrophages. Scaffold

strips that were kept unseeded (n=5) were also placed in 2 ml vials containing the same

medium, but without cells. The vials were rotated for 16 hours in a humidified atmosphere

containing 5% CO2 at 37°C to allow cells adhere to the scaffold strips. After seeding, each

scaffold strip was placed into a well of a 12-wells plates containing 1.5 ml culture medium

and 50 ng/ml PMA for another 24 hours. In 2D cultures, 1.25*106 cells were plated into 6-

wells together with 2 ml culture medium and 2 µl PMA for 48 hours.

Figure 5.1 Experimental set-up of the 2D and 3D cultures. Monocytes were transformed into adhering macrophages by PMA. To obtain different cell phenotypes, different cytokines were added to the cell culture medium to allow for macrophage polarization, or cells were kept unpolarized. At day 10 and 20, scaffold strips were replenished to maintain a constant cell population on the fibers. 2D samples were sacrificed after 1, 3, 6, 8, and 10 days, and the 3D samples were sacrificed after 2, 7 and 28 days.

5.3.2.3 Polarization and culture of macrophages

Cells in 2D cultures and cells on seeded scaffold strips were polarized into the

inflammatory or healing type of macrophages referred to as M1 or M2, respectively, or

kept unpolarized, referred to as M0 (n=24 per macrophage type). M1 polarization medium

consisted of 100 ng/ml Lipopolysaccharide S ((LPS), Sigma Aldrich) and 20 ng/ml

Interferon gamma ((IFN-ɣ, Peprotech, London, UK) in culture medium. M2 polarization

medium consisted of 20 ng/ml of Interleukin-4 (IL-4) and IL-13 (Peprotech, London, UK) in

culture medium. M0 cells were cultured using culture medium without additional

cytokines. Polarization and medium change was performed 3 times per week. Medium

was stored at -80°C until further use. 2D cell cultures were kept in culture for 1, 3, 6, 8,

and 10 days (n=2 per macrophage phenotype per time point), and photos of the cell

morphology were taken before sacrifice, using the moticam 2500 camera (Motic AE2000

Trino, Wetzlar, Germany). Seeded scaffold strips were cultured for 2 days, 1 week and 4

weeks, (n=8 per macrophage phenotype per time point). Unseeded scaffold strips were

Chapter 5

78

kept in culture for 2 days (n=1), 1 week (n=2) and 4 weeks (n=2). To maintain cells on the

scaffolds throughout the culture period and as macrophages are known to migrate out of

the scaffold or to undergo apoptosis, the samples of the 4 week culture groups were

replenished with cells. After 10 and 20 days, 1.0*106 and 1.5*106 cells, respectively, were

added to these scaffold strips, using the rotating seeding method, as described above.

After culture, the strips for SEM analysis (n=2 per phenotype per time point) were fixated

in formalin for 24 hours, which was replaced with sterile PBS afterwards and stored at 4°C

until analysis. Samples for DNA (n=3 per phenotype per time point) and gene expression

analysis (n=3 per phenotype per time point) were washed in sterile PBS and stored at -

80°C until further use.

5.3.3 Medium analyses

An esterase assay was performed to determine the esterase activity in culture medium

samples of 3D cultures (n=3 per time point and macrophage type) as a measure for the

amount of secreted enzymes by the macrophages. Esterase activity can be measured with

the use of p-nitrophenyl butyrate (pNPB, Sigma-Aldrich) as a substrate. When incubated

with esterase, the substrate is hydrolyzed and the yellow p-nitrophenyl (pNP) is released,

which can be measured by UV absorbance at 405 nm. 150 µl of medium sample was added

to a well of a 96-well plate, together with 10 µl pNPB solution (3.55µl pNPB in 5.0 ml

acetonitrile (Sigma-Aldrich)) and 140 µl Tris-Buffer. The plate was incubated at 37°C for

30 minutes, at 405 nm, using a microplate reader (Multiskan GO, Thermo Scientific) and

readings were performed every 5 minutes. Cholesterol esterase (C3766, Sigma-Aldrich)

was used as a positive control, and all samples were measured in duplicate. One unit of

esterase activity was defined as 1 nmol pNP released from pNPB per minute at 37°C.

5.3.4 DNA assay

The total amount of DNA was determined as an indicator of cell number during culture.

This was used to normalize the outcome of the esterase activity per amount of cells.

Scaffold strips were lyophilized after culture. Lyophilized tissue samples were digested in

300 µl papain solution (100 mM phosphate buffer (pH=6.5), 5 mM L-cysteine, 5 mM

ethylene-di-amine-tetra-acetic acid (EDTA), and 140 μg papain per ml, all from Sigma) at

50°C for 18 hours. After centrifuging the samples, the digest supernatant was collected

and used for the DNA assay. The amount of DNA in the samples was determined using the

Hoechst dye method [162] and a standard curve prepared of calf thymus DNA (Sigma

Aldrich). Using the assumption that all cells contain 6.5 pg of DNA [163], the amount of

cells per scaffold was calculated. In 3D cultures, DNA amounts were calculated per mg dry

weight scaffold. For 2D cultures, cell number was measured using the Tali® Image-based

cytometer (Invitrogen, Breda, The Netherlands). Measurements were averaged per

group.


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5.3.5 Morphology of scaffold fibers

Scaffold fibers were visualized by scanning electron microscope (SEM) to analyze the

morphology of the fibers. Formalin-fixed samples were stored in sterile PBS until use.

Samples were dehydrated in a graded ethanol series, starting from 50% to 100% in 5 to

20% increments. The ethanol was then allowed to evaporate, and samples were visualized

by SEM (Phenomworld, Eindhoven, The Netherlands). After SEM analysis, samples were

treated with 4.6% natrium hypochlorite for 15 minutes at room temperature, to remove

all cells. Samples were washed twice in water and the same samples were analyzed by

SEM again, in order to visualize parts of the scaffold fibers that were covered by the cells

during the first SEM analysis.

5.3.6 Gene expression analysis

The cultured scaffolds, stored at -80°C, were disrupted with a microdismembrator

(Sartorius, Goettingen, Germany) 3 times for 30 seconds at 3000 rpm, using RNA-se free

metal beads (diameter 3 mm, Sartorius, Goettingen, Germany) in Nalgene® cryovials

(Sigma-Aldrich) to homogenize the samples. Cells of all samples were lysed using RLT

buffer with β-mercaptoethanol (Sigma-Aldrich). RNA was isolated with Qiagen RNeasy kit

(Qiagen, Venlo, The Netherlands) according to the manufacturer’s protocol, including a

DNA-se incubation step. RNA quantity and purity were determined with a

spectrophotometer (NanoDrop®, ND-1000, Isogen Life Science, IJsselstein, The

Netherlands). Subsequently, cDNA was synthesized starting from 200 ng RNA in a 25 µl

reaction volume consisting of 20 ng/ml random primers (Promega, Madison, WI), 5 mM

dNTPs, 5x first strand buffer, 0.1M DTT, 200U/µl M-MLV Reverse Transcriptase (RT)

(Invitrogen, Breda, The Netherlands) and double autoclaved water (ddH2O). cDNA

synthesis was performed in a Thermal Cycler (C1000 Touch ™, Bio-Rad) by subjecting the

samples to a temperature cycle of 72°C for 6 minutes, 37°C for 5 minutes (with subsequent

addition of M-MLV), 37°C for 60 minutes, and 95°C for 5 minutes. Absence of genomic

contamination was checked using PCR (iCycler iQ, Bio-Rad, Hercules, US) with

glyceraldehyde-3-phosphate dehydrogenase (GAPDH) primers.

Gene expression levels of genes expressed by inflammatory-type macrophages (TNF-α,

CCR7, MCP-1 and IL-23) and healing-type macrophages (TGF-β1, VEGFA, MMP9, IL-10,

CD163, and MRC-1) were measured, and GAPDH was selected as a reference gene. The

primer sequences are listed in Table 1. Gene expression levels were determined by adding

20mM primer mix to the cDNA templates, together with SYBR® Green Supermix (Bio-Rad)

and ddH2O. All samples were analyzed in duplicates. The real-time PCR reaction was carried

out for 3 minutes at 95°C, 40x (20 seconds at 95°C, 20 seconds at 60°C, 30 seconds at

72°C), 1 minute at 65°C, followed by a melting curve analysis (iCycler MyiQ, Bio-Rad,

Hercules, US). Ct values were normalized to the GAPDH reference gene and the expression

levels were calculated using the formula 2−ΔCt .

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80

Table 1 List of primer sequences used in gene expression analysis.

Primer Symbol Accession

number

Primer Sequence (‘5-‘3)

Tumor necrosis factor alfa TNF-α NM_000594 FW: GAGGCCAAGCCCTGGTATG

RV: CGGGCCGATTGATCTCAGC

Interleukin 23 IL-23 NM_016584 FW: AGCTTCATGCCTCCCTACTG

RV: CTGCTGAGTCTCCCAGTGGT

Chemokine (C-C motif) receptor 7 CCR7 NM_001838 FW: AAGCCTGGTTCCTCCCTATC

RV: ATGGTCTTGAGCCTCTTGAAATA

Monocyte chemoattractant protein 1 MCP-1 NM_002982 FW: CAGCCAGATGCAATCAATGCC

RV: TGGAATCCTGAACCCACTTCT

CD163 molecule CD163 NM_004244 FW: CACTATGAAGAAGCCAAAATTACCT

RV: AGAGAGAAGTCCGAATCACAGA

Transforming growth factor, β1 TGF-β1 NM_000660 FW: GCAACAATTCCTGGCGATACCTC

RV: AGTTCTTCTCCGTGGAGCTGAAG

Vascular endothelial growth factor A VEGFA NM_001025366 FW: GCAGAATCATCACGAAGTGG

RV: GCATGGTGATGTTGGACTCC

Interleukin 10 IL-10 NM_000572 FW: GACTTTAAGGGTTACCTGGGTTG

RV: TCACATGCGCCTTGATGTCTG

Matrix metalloproteinase 9 MMP-9 NM_004994 FW: TGGGGGGCAACTCGGC

RV: GGAATGATCTAAGCCCAG

Mannose receptor, C type 1 MRC-1 NM_002438 FW: TGGGTTCCTCTCTGGTTTCC

RV: CAACATTTCTGAACAATCCTATCCA


Statistics were performed using GRAPHPAD Prism (version 5.04) and differences were

considered significant for p-values <0.05. All data are presented as mean ± standard error

of the mean. Regression analyses were performed to determine changes in amount of

DNA over time. In case of a significant in- or decrease with p<0.05 or p<0.01, the increase

or decrease was calculated from the predicted model equation and expressed as

percentage. In 3D cultures, we assumed a one-sided population, as only a decrease in

amount of DNA was expected. Statistical differences in qPCR data were analyzed with one-

way ANOVA followed by a Tukey’s multiple comparison post-hoc test. The LPS/IFNɣ

polarized samples after 1 week and 4 weeks of culture were not included in statistical

analysis of gene expression, due to the limited amount of RNA extracted from these

samples.


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5.4 Results

5.4.1 Morphology and number of polarized macrophages in 2D cultures

Representative pictures of THP-1 monocytes in suspension and the adherent

macrophages, which were cultured in 5.2D and either left unpolarized or polarized with

LPS/IFNɣ or IL4/IL13, are shown in Figure 5.2A. Cell diameter increased when monocytes

adhered to the culture flasks. When cells were polarized with LPS/IFNɣ they became

elongated, while polarization with IL4/IL13 showed rounded cells with a larger cell

diameter compared to the unpolarized cells. DNA amount in the macrophage cultures

decreased with time for all phenotypes (Figure 5.2B), however more prominent in the

LPS/IFNɣ polarized cells, as a higher plateau level for IL4/IL13 polarized cells compared to

LPS/IFNɣ polarized cells was found (p<0.001). After 10 days of culture, only 11% and 38%

(both p<0.01) of the initial amount of the cells polarized with LPS/IFNɣ or IL4/IL13,

respectively, remained in the culture wells.

Figure 5.2 Representative pictures of THP-1 monocytes and activated macrophages (A). Macrophages were kept unpolarized, or polarized towards an M1 or M2 phenotype by addition of the LPS/IFNɣ cytokines, or IL4/IL13, respectively. Unpolarized cells demonstrated a rounded morphology. Bigger rounded cells were observed when polarized with IL4/IL13, while cells became mainly elongated by addition of LPS/IFNɣ. The amount of DNA was reduced in all groups, however more prominent in the LPS/IFNɣ treated cells (B).

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5.4.2 Number of polarized macrophages on PCL scaffolds

Due to replenishment of the scaffold strips after 10 and 20 days, the amount of DNA

during culture in the unpolarized and IL4/IL13 polarized samples remained constant

(Figure 5.3). However, decreased amounts of DNA were observed of macrophages on PCL

scaffolds when polarized with LPS/IFNɣ to 0.5% of the initial levels after 4 weeks (p<0.05).

Figure 5.3 The amount of DNA during culture on PCL scaffold fibers. As a result on re-seeding, the amount of cells on the fibers remained stable in the unpolarized and IL4/IL13 polarized groups. A significant decrease in amount of cells was observed in LPS/IFNɣ polarized cells. Stars indicate time points of macrophage replenishment.

5.4.3 Morphology of polarized macrophages on PCL scaffolds

Two days after seeding, large populations of cells were observed on the scaffold, in each

group (Figure 5.4A, D, G). LPS/IFNɣ polarized cells appeared to be more spiky and showed

a rougher surface compared to the unpolarized and IL4/IL13 polarized cells, which showed

a rounded morphology. After 1 week, many LPS/IFNɣ polarized cells lost viability as we

observed many cell remnants on the scaffold fibers. Furthermore, the remaining cells

were smaller (Figure 5.4E) compared to the unpolarized (Figure 5.4B) and IL4/IL13 (Figure

5.4H) polarized macrophages. Both elongated and rounded cells were observed on the

scaffold fibers of the unpolarized and IL4/IL13 polarized groups. After 4 weeks, less cells

were found in the LPS/IFNɣ polarized samples, while many cells, single or in groups were

found in the unpolarized and IL4/IL13 polarized sample groups (Figure 5.4C, F, I).


83

5 Figure 5.4 Representative SEM images of cells cultured on scaffold meshes for 2 days, 1 week, and 4 weeks, which were unpolarized (A-C), polarized with LPS/IFNɣ (D-F), or polarized with IL4/IL13 (G-I). Decreasing amounts of cells were detected in time. No clear differences in cell morphology could be observed between the unpolarized and IL4/IL13 polarized cells, while the LPS/IFNɣ polarized cells showed to have a spiky appearance (D) with shrinkage of these cells after 1 week (E). White scale bar represents 50 µm.

5.4.4 Degradation of PCL scaffolds by polarized macrophages

Figures 5.5A-I show representative SEM images of scaffold fibers after removal of the

cells. After 2 days, unpolarized cells did not appear to degrade the scaffold fibers, as no

visual damage was observed in SEM images. Scaffold fibers of the IL4/IL13 polarized

samples showed minor damage in few spots, while this was more pronounced for the

LPS/IFNɣ polarized samples. Scaffolds containing the LPS/IFNɣ polarized cells showed the

highest degree of degradation of the scaffold fibers after 1 week of culture. Surface

erosion of the fibers was clearly visible at various spots in the scaffold strips, while this

was observed at only a few spots in the unpolarized and IL4/IL13 polarized groups. At 4

weeks, scaffolds showed similar results compared to the scaffolds after 1 week of culture.

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Up to 1 week of culture, no damage in terms of broken fibers, surface erosion or cracks in

the fibers were observed in the scaffold cultured without cells, while after 4 weeks, minor

surface erosion could be observed at some places in the scaffolds (Figure 5.5 J-L).

Figure 5.5 Representative SEM images of scaffold fibers after removal of cells. Photos were taken after 2 days, 1 week, and 4 weeks of culture with unpolarized cells (A-C), cells that were polarized with LPS/IFNɣ (D-F), polarized with IL4/IL13 (G-I), or scaffolds that were kept in culture without the presence of cells (J-L). All groups showed scaffold degradation, although this was most pronounced in the LPS/IFNɣ polarized group. When no cells were added to the scaffold, only minor damage was observed, which was due to hydrolytic degradation. White particles are natrium hypochlorite residues. White scale bar represents 30 µm.


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5

5.4.5 Medium analysis of 2D and 3D cultures of polarized macrophages

In order to determine the levels of secreted enzymes by the different macrophage

phenotypes, esterase activity assays were performed on culture medium samples of 3D

cultures. All phenotypes showed to be able to release esterases. When normalized to DNA

amount (Figure 5.6), enzyme activity in LPS/IFNɣ polarized cells was higher compared to

unpolarized and IL4/IL13 polarized samples when cells were cultured in 3D for 1 week

(P<0.05) and 4 weeks (p<0.001). Furthermore, esterase activity per DNA increased after 4

weeks of culture compared to 1 week in the LPS/IFNɣ polarized groups (p<0.001).

Figure 5.6 Esterase activity of the cells, corrected for the amount of DNA present in the scaffold meshes during culture. LPS/IFNɣ treated cells resulted to have increased esterase activity per cell compared to the other groups.

5.4.6 Gene expression analysis in 3D cultures

Expression of macrophage phenotypic and immune response genes was analyzed to

investigate whether a different set of markers was expressed when cells were cultured in

PCL scaffolds in unpolarized state or polarized towards the M1 or M2 phenotype (Figure

5.7). TNF-α levels were comparable between the groups after 2 days. After 1 week, the

expression levels of the unpolarized and IL4/IL13 polarized groups decreased compared

to the LPS/IFNɣ polarized group at 2 days (p<0.05). CCR7 and MCP-1 were significant

increased in the LPS/IFNɣ polarized groups after 2 days, compared to the unpolarized and

IL4/IL13 polarized groups (p<0.001). In time, CCR7 levels decreased in both the

unpolarized and IL4/IL13 polarized groups, with p<0.05 after 1 week for both groups, and

p<0.05 after 4 weeks in the unpolarized group. Expression levels of CD163 after 1 and 4

weeks increased in the unpolarized samples compared to the LPS/IFNɣ polarized samples

(p<0.01 and p<0.05, respectively). Furthermore, CD163 expression levels of the

unpolarized cells were increased compared to the levels found in the IL4/IL13 polarized

cells, after 1 week (p<0.05). MMP9 expression increased in time for the unpolarized

group, with higher levels found after 4 weeks, compared to 2 days and 1 week (p<0.05).

At this time point, MMP9 levels were also higher in comparison with the LPS/IFNɣ

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polarized group (p<0.001) and the IL4/IL13 polarized samples after 4 weeks (p<0.01).

Apart from expression levels of CD163 and MMP9 no other differences in expression

levels were observed between the unpolarized and IL4/IL13 polarized samples. No

differences in expression levels of IL-23, TGF-β, VEGF A, IL-10, and MRC-1 were observed

between the groups (data not shown).

Figure 5.7 Gene expression analysis of cells cultured on PCL scaffolds for 2 days (all groups), 1 and 4 weeks (unpolarized and IL4/IL13 polarized cells only). CCR7 and MCP-1 levels were increased in LPS/IFNɣ polarized cells compared to other groups. Besides minor differences between expression levels of CD163 and MMP9 after 1 and 4 weeks, respectively, no differences in expression levels were observed between the unpolarized and IL4/IL13 polarized samples. No differences in expression levels of IL-23, TGF-β, VEGF A, IL-10, and MRC-1 were observed between the groups (data not shown). # denotes significant differences compared to 1 week within the same phenotype with p<0.05, ^ denotes significant differences compared to 4 weeks within the same phenotype with p<0.05, while *, ** and *** denote significances of differences compared to the M1 phenotype after 2 days with p<0.05 and p<0.01, p<0.001.


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5.5 Discussion

Macrophages are claimed to play an essential role in tissue outcome and show a

continuum of functional properties alternating between the pro-inflammatory (M1) and

the tissue-healing (M2) phenotypes. In order to elucidate the effect of macrophage

phenotype on scaffold degradation 2D and 3D in vitro studies were performed.

During 2D cultures, we observed differences in cell morphology between the groups,

which is indicative for different phenotypes. Mainly elongated cells were observed

throughout the whole culture time when macrophages were polarized using LPS/IFNɣ and

mainly rounded cells in case of IL-4 and IL-13 treatment. This was in line with findings by

others [242, 244], while McWhorther et. al. [245] found the opposite morphology during

polarization of macrophages after 1 day. This might be due to the use of another cell

source and species by McWhorther, as they cultured mouse bone marrow derived

macrophages instead of human cells. Apparently, macrophages are very sensitive cells

and do not only react to cytokines present in their environment, but also to other

substances, such as the PMA concentration that influenced cell shape [246]. Furthermore,

it has been described by others that macrophage morphology is also surface-dependent

[247, 248]. This indicates that morphology may not be a reliable indicator of macrophage

phenotype and phenotype-specific markers should be used in addition to distinguish

between cell phenotypes. Also, to make a fair comparison between 2D and 3D cultures,

2D substrates should be made of the same biomaterial that is used in 3D cultures. This

was not performed within this study, as this is beyond the scope of our study.

Several studies demonstrated that macrophages were not only polarized by secreted

cytokines, as also the scaffold surface, cell shape, fiber diameter, pore size, and strain are

reported to influence the macrophage phenotype [140-145, 249]. This indicates that

scaffold morphology and composition can be used to promote an optimal healing

response. Garg et. al. showed that a high fiber diameter (>3 µm) together with a high

porosity (>80%) results in a transition of macrophages towards the M2 phenotype [144].

Also in our study, the results of qPCR data indicated a shift from the unpolarized

macrophages towards the M2 phenotype when exposed to the PCL scaffold, as apart from

expression levels of the CD163 and MMP9 genes, which were increased at some time

points for the unpolarized samples, no other differences in expression levels were

observed between the unpolarized and the IL4/IL13 polarized samples. This phenotypic

shift towards the M2 type is likely due to the scaffold morphology and composition. M2

macrophages can be further classified into different sub-phenotypes [136, 233]. The

differences in MMP9 and CD163 levels between the unpolarized cells and the IL4/IL13

polarized cells might be indicative for a shift of the unpolarized cells towards another sub-

phenotype of the M2 macrophages as compared to the sub-phenotype of the IL4/IL13

polarized cells. IL4/IL13 are believed to polarize the macrophages towards an M2a

phenotype, and MMP9 secretion is mainly observed in the M2b phenotype [233].

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Within our study, it seems that addition of cytokines overruled the effect of high fiber

diameter and porosity (10 µm and 80% respectively) which guided the unpolarized cells

towards the M2 phenotype, as after each burst of LPS/IFNɣ cytokines the macrophages

seemed to remain their M1 phenotype, and did not transdifferentiate towards the M2

phenotype, based on the amount of DNA over time and the qPCR data after 2 days.

The LPS/IFNɣ polarized cells showed a reduced viability in both 2D and 3D cultures

compared to unpolarized cells and IL4/IL13 polarized cells. This might be due to the

observed increased levels of TNF-α expression in the LPS/IFNɣ polarized cells, which is

known to induce cell death [250, 251]. This was in line with findings by others that showed

increased levels of TNF-α production in LPS/IFNɣ polarized cells [242, 252]. In a review by

Italiani et. al. [253] it was described that M1 cells are end-stage killer cells, which die

during the inflammatory response, probably due to its own nitric oxide (NO) production

[254]. Predominantly M2 type of macrophages were observed around implants in several

studies [95, 255]. However, it needs to be further clarified whether this is a transition from

the M1 phenotype towards the M2 phenotype within the same cells, within the cell

population, or whether a selective death of M1 macrophages due to apoptosis results in

a relative increase of cells with the M2 phenotype.

SEM images of scaffold fibers after removal of the cells showed more local damage of the

scaffold fibers in the LPS/IFNɣ polarized cells group compared to the unpolarized cells and

IL4/IL13 polarized cell groups. This was also expected, as LPS/IFNɣ polarized cells, are

known to produce high levels of enzymes and ROS that contribute to scaffold degradation

[136]. Esterase assays confirmed that the amount of secreted esterase per cell was

increased in the LPS/IFNɣ polarized cells compared to the unpolarized cells and IL4/IL13

polarized cells.

5.6 Conclusion

We were able to generate distinctive macrophage phenotypes in both 2D and 3D cultures

by the addition of cytokines, and observed that the life-span of LPS/IFNɣ polarized cells

was shorter compared to IL4/IL13 stimulated cells. In 2D, cell morphology was different

between the phenotypes, while this was less pronounced in the 3D cultures. Using qPCR

we were able to distinguish between the cell phenotypes present on the scaffold fibers.

Unpolarized macrophages in PCL scaffolds expressed similar genes as compared to the

IL4/IL13 polarized cells, which is indicative for a transition towards a M2 phenotype, and

is probably induced by the electrospun mesh, which is a beneficial feature for in situ tissue

engineering. We observed that all macrophage phenotypes were able to secrete

esterases, which are known to degrade the PCL scaffold fibers. All phenotypes indeed

contributed to local scaffold degradation, although this was far more pronounced in the

LPS/IFNɣ polarized cells compared the unpolarized and IL4/IL13 polarized cells. Given this


89

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knowledge, we suggest a correlation between macrophage phenotype and scaffold

degradation, with inflammatory macrophages accelerating degradation.

Acknowledgements


Institute for Regenerative Medicine (NIRM, grant No. FES0908). The authors gratefully

thank Roel Lalieu from Xeltis for electrospinning of the PCL scaffolds. Virginia Ballotta is

acknowledged for providing some of the primers.

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General discussion

6

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Worldwide, cardiovascular diseases are the major cause of death, resulting in a need for

800.000 bypass surgeries and 290.000 heart valve replacements each year [11, 22, 256].

With the ageing population and improved diagnostic methods, this number will only

increase. The currently available replacements use non-living materials, and are as a

consequence unable to remodel and adapt to changes in their environment. Tissue

engineering attempts to overcome these shortcomings, and aims at implantation of

bioresorbable grafts that will fully integrate with the host tissue, leaving behind a living

tissue that is able to adapt and remodel. Various tissue engineering approaches have been

investigated in the past with the intention of effective regeneration of diseased

cardiovascular tissues. In situ tissue engineering offers several advantages over the

classical in vitro tissue engineering approach, as this approach aims to employ off-the-

shelf products, reducing the costs and time to produce replacements, and comprises less

regulatory issues. The main challenge is to find the proper biomaterial that can be used

to create matrices that maintain good mechanical integrity, immediately after

implantation and start resorbing when sufficient neo-tissue has been formed that can take

over this role. As different properties are desired in various applications, it is of high

importance to be able to tune the mechanical properties and resorption characteristics of

biomaterials. The focus of this thesis was to investigate degradation characteristics of

electrospun scaffolds, manufactured from different supramolecular biomaterials.

Furthermore, the interplay between scaffold degradation rate and the amount and

composition of neo-tissue was examined.

6.1 Main findings of the thesis

6.1.1 Slow-degrading scaffold material reduces compaction and retraction

A balance between tissue formation and scaffold degradation is believed to be important

to maintain a functional replacement with proper mechanical integrity. Slow tissue

synthesis or too fast resorption of the implant might disturb this delicate balance, and has

shown to result in compaction and retraction of in vitro tissue engineered heart valves,

causing regurgitation in vivo [71, 72, 257]. Traction forces exerted by smooth muscle cells,

the cells that lay down a new layer of tissue, are believed to enhance this process [74,

151]. During tissue remodeling, the amount of these cells will decrease, which therefore

probably also results in lower traction forces. To bridge the phase where tissue is not

mature enough to withstand the forces exerted by the cells and constant loading of the

implant, a scaffold with sufficient mechanical integrity during the first months after

implantation is desired. Therefore, we studied in vitro whether the use of slow- (PCL)

instead of a fast-degrading (PGA-P4HB) electrospun scaffold meshes, and a lower cell

passage number to enhance tissue formation, reduces compaction (chapter 2). Reported

time for complete in vivo bioresorption of PGA-P4HB varies from 4 (PGA) to 8 (P4HB)

weeks [68] to 4-6 months for PGA-P4HB, with 50% loss of mechanical properties within 2

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6

weeks [258]. Complete bioresorption of electrospun PCL scaffolds is reported to be at

least 1-2 years [58, 183]. We demonstrated that the use of a slow-degrading material

resulted in improved resistance to retraction of tissue engineered valvular leaflets and

reduced compaction of strips compared to fast-degrading material. Tissue formation,

stiffness and strength increased with decreasing passage number, however, this did not

influence compaction. Furthermore, tissue constructs were engineered using both ovine

and human cells, to determine the effect of interspecies differences on tissue

development. Although variations between the actual amount of ECM components were

found between the species, the effects e.g. on compaction were comparable. Overall, in

terms of compaction, the influence of the scaffold type seemed larger than the influence

of the tissue production of several cell sources.

6.1.2 Organized tissues which maintain their 3D shape when cultured onto slow-

degrading scaffold materials contain similar ECM values compared to native pulmonary

heart valves

Scaffolds with slow- and fast-degradation rates will contribute to the mechanical integrity

of the implant differently with time. We hypothesized that cells on fast-degrading scaffold

material will produce increased amounts of tissue compared to cells on slow-degrading

materials, to compensate for the loss in mechanical integrity. Using in vitro tissue

engineering, we studied tissue evolution, in terms of ECM composition and mechanical

properties of the constructs in time, of vascular cells cultured on slow- (PCL) or fast-

degrading (PGA-P4HB) electrospun scaffolds (chapter 3). It was shown that tissues

cultured on slow-degrading scaffolds contained organized tissue formation maintaining

their 3D shape during culture, while the tissues cultured on fast-degrading scaffold

materials demonstrated appositional growth and compaction during culture. This again

demonstrated that slow-degrading scaffold material is favored over fast degrading

scaffolds to ensure stable mechanical integrity during the initial phase after implantation.

During the first two weeks of culture, when the PGA-P4HB scaffold was clearly degrading,

the cells cultured onto these scaffold meshes were more synthetic in agreement with our

hypothesis. However, this synthetic phenotype was only a temporary feature, as after 6

weeks lower amounts of sGAG and collagen were measured in the PGA-P4HB based

tissues when compared to the PCL-based tissues. Compaction of the PGA-P4HB scaffolds

resulted in a smaller surface area, and less volume for the cells within these tissues to lay

down their ECM, compared to PCL-based tissues, which might have resulted in this higher

amount of ECM after 6 weeks of culture in the latter. To make a fair comparison between

tissue composition of the in vitro engineered tissues grown on scaffolds with a different

degradation rate, and between in vitro engineered tissues and native tissues, we

described a method to correct for the amount of remaining scaffold weight. Implementing

this correction, we found ECM values that were similar to, or towards values of native

pulmonary heart valves. Although collagen crosslink values were increasing with in vitro

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culture time, the measured values were still lower in engineered tissues compared to their

native counterparts.

6.1.3 PCL and PCL-based supramolecular polymers exhibit different degradation

characteristics

As in vitro culture of cardiovascular substitutes is expensive and time-consuming, in situ

tissue engineering would be a good alternative. This would also allow for off-the-shelf

availability of implants with reduced production time and less regulatory issues that are

related to tissue culture. Unpublished data by our group demonstrated that the slow-

degrading PCL material resulted in plastic deformation when it was cyclically loaded.

Consequently, this material is not suitable for in situ tissue engineering of heart valves, as

dynamic loading of the leaflets places high demands on the scaffolds immediately after

implantation. Supramolecular biomaterials contain hydrogen bonding motifs, like UPy or

BU incorporated into their molecular structure, resulting in suitable materials for in situ

heart valve tissue engineering.

In vivo, resorption of the implanted material can be accomplished via two main pathways,

which includes the enzymatically accelerated hydrolytic pathway and the oxidative

pathway. To investigate both pathways, separately and in an accelerated fashion, in vitro

degradation assays were designed. With the use of these models, degradation

characteristics of several promising supramolecular materials were explored and

compared to the conventional PCL material (chapter 4). We demonstrated that,

depending on the morphological and chemical composition of the materials, conventional

and supramolecular PCL-based electrospun meshes responded differently to both

pathways. The enzymatic accelerated hydrolytic pathway mainly affected conventional

PCL scaffolds, while supramolecular materials were not (PCL-UPy) or only mildly (PCL-BU)

affected, which was enzyme dependent. PCL material was not susceptible to oxidative

degradation. The supramolecular PCL-UPy materials were, dependent on the PCL soft

segment length and the supramolecular moiety coupled to the PCL backbone, not

susceptible (PCL800-UPy) or susceptible (PCL2000-UPy and PCL-BU) to oxidative

degradation. When materials were treated with the enzymatic solution, surface

degradation seemed to be the dominant degradation mechanism for PCL and PCL-BU.

When exposed to oxidative solutions, surface erosion was observed for PCL-BU, while

both surface erosion and bulk erosion was seen in the PCL-UPy materials.

This insight into the degradation characteristics of PCL-based (supramolecular) materials

allows us to tailor degradation characteristics. Different combinations of polymer

backbones modified with supramolecular moieties can be created which result in various

polymer properties such as mechanical properties and/or resorption rate. This mix-and-

General discussion

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match toolbox can be utilized to screen and select the relevant biomaterial for pre-clinical

in vivo studies targeted to different clinical applications.

6.1.4 Electrospun PCL scaffolds guide tissue regeneration and are mainly degraded by

inflammatory macrophages

During the immune response, which starts immediately after implantation of a scaffold

material, macrophages are known to play an important role in the resorption of the

implant. Macrophages possess plastic functional properties and represent a continuum in

which they can alternate, dependent on micro-environmental factors, between the pro-

inflammatory (M1) and the healing (M2) macrophages [136, 234]. Polarization of

macrophages towards the healing-type induced by the scaffold material is desired to

improve final tissue outcome. To illustrate whether there is a correlation between the

inflammatory or the healing macrophage phenotypes and the degradation of electrospun

meshes, in vitro tests were developed and used (chapter 5). In 2D culture, a clear

morphological difference was observed when different cytokines were added to the

macrophages in order to polarize them towards the inflammatory or healing phenotype,

which was less pronounced in 3D cultures. Cells of the inflammatory phenotype had a

shorter life-span compared to cells of the healing phenotype or the untreated

macrophages, in both 2D and 3D cultures. During 3D cultures of different macrophage

phenotypes, we visually observed enhanced degradation of the scaffold fibers by the

inflammatory phenotype. No difference, in terms of scaffold degradation, was observed

between the healing phenotype and the unpolarized macrophages, which implies that the

latter polarized towards the healing phenotype in 3D cultures. This was confirmed by gene

expression studies, where the gene expression profile of the unpolarized macrophages

was most similar to the gene expression profile of the healing macrophages.

In summary, the results presented in this thesis suggest that the choice of a scaffold

material is of high importance to maintain a good balance between scaffold degradation

and tissue formation, and therewith maintaining mechanical integrity. A slow-degrading

material is favored over a fast-degrading material, as mechanical integrity will be

maintained for a longer period, which is mainly important in in situ tissue engineering

purposes where a bare scaffold is implanted. Furthermore, neo-tissue seemed to be

better organized when cultured on slow-degrading scaffold materials and less prone to

compaction. Macrophages are known to play an essential role in scaffold degradation.

Although all macrophage phenotypes seemed to be able to degrade scaffold material in

vitro, we visually observed a higher degree of scaffold degradation by the inflammatory

phenotype compared to the healing phenotype. Untreated macrophages that were

cultured into the scaffold, polarized into the healing phenotype, indicating that the used

material is guiding tissue regeneration rather than repair.

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6.2 Towards the most promising tissue engineering approach and

scaffold material

6.2.1 In vitro tissue engineering versus in situ tissue engineering

In order to replace diseased cardiovascular tissues, different tissue engineering

approaches are explored, as described in the introduction of this thesis. Although

promising results are obtained within every approach, they also all have their challenges

that require further optimization before safe translation to the clinic. The classical in vitro

tissue engineering approach requires months before an engineered construct is suitable

for implantation. Decellularization of the in vitro tissue engineered constructs improves

the classical approach, in terms of readily available, off-the-shelf products. This

decellularization approach also provides tissues replacements for those patients who do

not have tissues available for cell isolation, or are in such a critical situation that a waiting

time of a few months would induce high mortality risks. Despite this improvement, the

decellularization approach has to overcome more challenges before it will result in

reliable implants, as compaction and retraction of decellularized in vitro tissue engineered

heart valves is still a hurdle to overcome [43, 71-73]. Results of this thesis show that the

use of slow-degrading scaffold materials probably will result in further improvements.

Furthermore, changes in geometry could help in preventing compaction and retraction.

In situ tissue engineering using synthetic materials is a promising approach. It

demonstrates off-the-shelf availability, while production-time and -costs and regulatory

issues are significantly reduced compared to classical in vitro tissue engineering methods.

The in situ tissue engineering approach relies on the regenerative capacity of patients. A

challenge that needs to be addressed is the search for the appropriate scaffold material

in relation to the specific application. An immune response needs to be triggered and

controlled after implantation to ensure migration of different cell types towards the

implanted material. These include macrophages, which are involved in resorption of the

scaffold material and regulation of the healing response, and cells that generate the neo-

tissue. Several groups aim at active cell capture of specific cell-types involved in tissue

formation. Therefore, they create bioactive scaffold fibers, by incorporating peptides to

the material, or coat the fibers with growth factors or antibodies, which are released after

implantation in order to recruit specific cells [259, 260].

When certain cells are actively captured in the scaffold material, the natural immune

response is influenced and might even be disturbed. Whether the bioactive scaffold

materials have a beneficial or negative effect on the final tissue outcome needs to be

further elucidated. In situ tissue engineering without the addition of peptides, cytokines

or cells to the scaffolds relies on the properties of the material together with a natural

healing response and endogenous tissue growth (ETG) by the body itself. This is beneficial,

because the absence of bioactive moieties reduces regulatory requirements and simplifies

General discussion

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the production process as only synthetic biomaterials are involved. Promising pre-clinical

[90, 91] and clinical data [92] were gathered using this approach.

6.2.2 Choosing the appropriate scaffold material

Electrospun scaffold materials aimed to replace cardiovascular tissues by the in situ

approach need to fulfill specific requirements [261]. They should be biocompatible, to

prevent a severe inflammatory reaction that might impair the healing cascade, or even

cause rejection by the body. Another requirement is controlled resorption of the material,

with non-toxic resorption products that are removed from the body without side-effects.

Dependent on the application, higher or lower mechanical forces will be applied to the

material. It is of high importance that the scaffold is able to withstand these forces and

ensures mechanical integrity immediately after implantation, until neo-tissue is able to

take over this function. The supramolecular materials PCL-BU and PCL-UPy are promising

materials to be used in in situ tissue engineering, as they contain strong and elastic

properties that can withstand the repetitive loading forces that are exerted immediately

after implantation, when the scaffold is implanted as a heart valve replacement. Cells

behave differently when growing on a stiff or elastic material. For example, PCL is known

to be stiffer compared to PCL-BU. Therefore, the PCL-BU scaffolds display more stretch in

vivo compared to PCL scaffolds when implanted at the same anatomical location.

Scaffold architecture also influences tissue outcome and additional requirements can be

added to promote optimal healing responses. First, scaffolds require an interconnected

pore structure with a high porosity to allow for cell infiltration and tissue formation,

including vascularization [262, 263]. In addition, it has been shown that fiber diameter not

only influences porosity and, thereby, cell infiltration, but also influences the polarization

of macrophages during the inflammatory response. A healing phenotype was observed in

scaffolds with a high fiber diameter (>3 µm) together with a high porosity (>80%) [144,

145]. This was also observed within our experiments in chapter 5. Macrophages

preferentially polarized towards the healing type when cultured on PCL scaffolds with a

fiber diameter of 10 µm and a porosity of 80%. Furthermore, it is demonstrated that

alignment of the scaffold fibers stimulates the cells to increase collagen synthesis [264,

265].

Taken together, we aim for a scaffold that allows cell infiltration, modulates the immune

response, and supports tissue formation that is able to remodel and gradually takes over

the mechanical functionality of the scaffold. Ideally, the scaffold should be completely

resorbed, to prevent ongoing inflammatory responses against the foreign body material.

Finally, easy surgical handling is an additional benefit for clinicians who are implanting

these scaffolds.

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6.2.3 Focus on in vivo resorption characteristics

6.2.3.1 Balance between scaffold resorption and tissue formation

An important aspect to consider when choosing a biomaterial is its resorption

characteristic. In case of too slow resorption of the material stress shielding of the neo-

tissue might occur, thereby impeding the regeneration process [231]. Furthermore, slow

resorption might lead to other undesirable outcomes, like a prolonged inflammatory

response. When the resorption process is too fast the mechanical integrity of the implant

is not maintained, as the neo-tissue is not sufficiently developed yet to bear the full

mechanical force required [232], leading to failure of the implant. The use of fast-

resorbing materials might be one of the reasons that contributed to heart valve leaflet

retraction and compaction observed in several in vivo studies [43, 71-73]. This is a result

of traction forces exerted by αSMA positive cells, likely in combination with an imbalance

of the newly formed tissue and fast loss of mechanical integrity of the scaffold due to

resorption [74, 148, 151]. As αSMA is related to traction forces of the cells [152], and

αSMA positive cells were demonstrated to decrease again in vivo [6], these traction forces

will also be decreased. Therefore, a slow-resorbing scaffold with sufficient mechanical

integrity during the first phase after implantation is desired to withstand the cell traction

forces during this period. Furthermore, we demonstrated in chapter 3 of this thesis that

in vitro less organized tissue was formed, that lost its original shape when cultured onto

fast-degrading scaffold materials. Findings by Hasizume et. al. [266] also demonstrated

that implantation of a slow-resorbing scaffold patch to treat chronic ischemic

cardiomyopathy in rats resulted in beneficial results in terms of cardiac function and

histology compared to faster resorbing patches. In addition, de Jonge et al. observed

during an in vitro study that after 2 weeks of culture, the newly formed collagen fibers

were not dense enough yet to withstand the traction forces of the cells and resulted in

collagen reorientation [49]. This suggests that a slow-resorbing scaffold material should

be chosen, that allows the newly formed collagen fibers to mature first, and therewith are

able to withstand the loads that are applied on the constructs, when aiming at maintaining

collagen orientation.

6.2.3.2 Size and anatomical location

The size of the grafts might also influence the material selection. Both cells from the

bloodstream and the adjacent tissue site will infiltrate into the graft, followed by tissue

formation. Complete cell infiltration and corresponding tissue formation throughout the

graft is expected to occur faster in short replacements e.g. short interpositional vascular

grafts, where both transmural and transanastomotic ingrowth contribute to fast cell

infiltration throughout the replacement, compared to long replacements like

reconstructions of aortic aneurysms where a larger area needs to be infiltrated. This

would suggest the use of a slow-resorbing material when long tissues need to be replaced.

In addition, the anatomical location of the implant might influence both the cell

General discussion

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infiltration and the resorption rate of a biomaterial. Mechanical forces, like compression,

fatigue and shear stress, or other factors like flow, pH and the presence of enzymes might

affect the resorption rate of the implanted material [121].

With the use of slow-resorbing materials, the question might arise what will happen with

the neo-tissue when the scaffold mesh is completely resorbed by the body. We

hypothesize that the neo-tissue has prolonged time to mature when growing on slow-

resorbing scaffolds and, thereby, is able to withstand the loading forces applied on the

constructs. Although preliminary results [103] showed decreased mechanical contribution

of the implanted heart valve scaffolds in the ovine model after 6 months, longer-term in

vivo experiments are needed to demonstrate the fate of the tissue after complete loss of

mechanical integrity of the scaffolds. Long-term animal experiments up to when the

implant is completely resorbed, should be performed to investigate the final outcome.

6.2.3.3 The interplay between direct cell contact and resorption of the scaffold fibers

It is suggested that direct contact between the cells and the scaffold fibers is needed in

order to degrade the scaffold material [131, 267]. We also found indications that are in

line with this hypothesis as pilot in vitro studies on the interplay between macrophages

and scaffold fibers showed surface erosion of the scaffold fibers by the macrophages,

clearly visible when cells were removed from the scaffold fibers (Figure 6.1). No surface

erosion or cracks were observed in fibers of scaffold meshes that were cultured in the

same medium, and thereby are in contact with the same amounts of enzymes and ROS

products, however separated from the macrophages with a porous membrane to avoid

direct cell contact. This indicates that direct cell contact is needed, probably to very locally

create a high concentration of enzymes and/or ROS products to degrade the scaffold

material.

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Furthermore, in vivo pilot studies using the ovine model [103] demonstrated mainly

scaffold resorption at those sites where macrophages infiltrated into the porous scaffolds

and tissue was formed (Figure 6.2). Infiltration of cells and tissue formation in the valves

was observed to start from the wall, and with time cell infiltration further towards the

leaflet tips was observed. This is in line with previous findings where implantation of

decellularized in vitro tissue engineered heart valves demonstrated fastest repopulation

with highest densities in the tissue wall compared to the heart valve leaflets [73]. The

observation that tissue formation always precedes resorption, even though local

differences are observed in rates of regeneration is an important safety aspect related to

in situ implantation of scaffold materials.

Figure 6.1 Schematic figure of a transwell assay experiment (A). In the upper compartment, cells are in direct

contact with the scaffold material. The secreted enzymes and oxygen radicals can migrate through a porous

membrane towards the lower compartment, where they can reach the scaffold material that is not in direct

contact with the cells. SEM images were taken from the scaffold in the upper compartment with cells (B), and

after removal of the cells (C). Small holes and surface erosion (black arrows) were observed in the scaffolds that

were in direct contact with the cells. The damage was observed underneath the cells and in the near environment

(<5 µm) where cells adhered to the scaffold fibers. No holes or surface erosion due to cells could be observed in

the scaffold that was cultured in the lower compartment (D), although some general hydrolytic degradation (white

arrow) was observed which is common after 4 weeks of culture in an aqueous solution. White scale bars represent

20 µm.

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Figure 6.2 SEM images of a 6-month valvular explant [103]. Cell infiltration and corresponding tissue formation started from the wall and belly region of the leaflets (C). With time cells infiltrated in the center of the leaflets (B) and further towards the leaflet tips (A). After removal of the tissue, we observed that at those places where no or few cells were infiltrated and therewith no or little tissue was formed, scaffold fibers were only minimally affected (D). When many cells were infiltrated, which was correlated to a large amount of tissue formation, scaffold fibers were severely affected by the infiltrated cells (E). Black and white

scale bars represent 1 mm and 20 µm, respectively.

Taken together, this would advocate use of a slow-resorbing material to ensure sufficient

time for the neo-tissue to form, remodel and mature before the mechanical integrity of

the scaffold material is completely lost. As the desired resorption properties are

depending on the application, the need to tailor the properties of bioresorbable polymers

specifically to the intended application is necessary.

6.3 Study limitations and the future of in situ cardiovascular tissue

engineering

6.3.1 Benefits of in vitro models

In order to understand tissue development on implanted scaffold meshes and how this is

affected by scaffold resorption, in vitro model systems providing a 3D environment are

very useful. With these in vitro models, we can obtain insight in the interactions between

cells and scaffold materials and tissue formation. Furthermore, the necessity of animal

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experiments is reduced with in vitro models. In chapter 2 and 3, we created rectangular

electrospun scaffold strips that were attached to metal rings [268, 269]. This model

system was used as a representative of in vitro tissue engineered heart valves. It allows

for high-throughput analysis, while using less scaffold material and cells in comparison to

complete full-scale heart valves. Furthermore, the use of complex bioreactors can be

circumvented.

In addition, the in vivo environment is very complex with plenty of factors involved at the

same time. In vitro models allow to unravel research questions with better controlled

environmental conditions. We particularly made use of this in chapter 4, where the

scaffold degradation pathways were investigated separately, while in vivo they are

occurring at the same time and might influence each other. The in vitro study provided us

with useful insights into the degradation characteristics of different materials that could

not be obtained with in vivo studies.

6.3.2 Limitations of in vitro models

Although our in vitro models provide us with further insight in tissue development and

degradation characteristics, they also encounter some limitations. The geometry of the

rectangular strips is simplified compared to the complex structure of a heart valve and, as

these strips were attached to metal rings, they encounter different forces compared to

heart valves. After obtaining further insights with the model system using strips, heart

valves can be cultured, as was done in chapter 2, in order to confirm the findings in the

clinically relevant geometry.

As we cultured our strips statically in chapters 2, 3, and 5, the cells within the scaffold

strips did not encounter the effect of cyclic loading, strain and shear stress due to blood

flow, all known to have an effect on valve interstitial cell behavior and tissue formation

[270, 271]. Previous in vitro results have demonstrated increased amounts of tissue

formation when constructs were cultured under intermittent straining protocols

compared to static controls [272, 273]. Also immune cells do sense differences in strain,

as a study by Ballotta et al. showed that around 7% strain resulted in polarization of

macrophages towards the healing-phenotype [249]. Furthermore, it was also shown

within our group that shear stress influences the recruitment of monocyte subsets, which

is of importance in in situ tissue engineering [274].

To investigate the correlation between macrophage phenotype and scaffold degradation,

we made use of a human monocytic cell line. By adding PMA, these monocytes were

differentiated into macrophages. Thereafter, macrophages were polarized into different

phenotypes by the addition of cytokines to the cell culture medium. This cytokine-

supplemented-medium was only refreshed every 2-3 days to provide new cytokines and

nutrients to the cells, while in vivo there is continuous refreshment of nutrients and

cytokines. In addition, in vivo cell replenishment occurs all the time as monocytes can

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migrate towards the biomaterial surface continuously where they differentiate into

macrophages, while we re-seeded the scaffolds every 10 days only.

Taken all these aspects into account, in vitro we can only partially replicate all aspects of

the in vivo situation. Therefore, pre-clinical experiments will always remain of high

importance for validation of the approach before clinical application. In case of heart valve

implants, the ovine model is the model of choice for pre-clinical experiments [275], as it

serves as the worst-case scenario in terms of calcification, one of the most common failure

in the prosthetic heart valves (Hopkins 2009) and has similarities to the human heart in

terms of anatomy and physiology.

6.3.3 Inter-patient variation

Human cells were used in the experiments described in chapter 2 and 3. It is known that

due to biological variation, different absolute values will be obtained when using cells of

different patients. Variation in tissue properties are not only observed within in vitro

engineered tissues, but are also found in native tissues [188]. Although inter-patient

variation will lead to variations in absolute values of in vitro tissue engineered constructs,

the overall effects in terms of e.g. mechanical properties or tissue composition due to a

certain treatment can be expected to be similar, as also described by research performed

within our group [48, 269].

Underlying diseases and the age of the patient might affect tissue remodeling or scaffold

resorption and, therewith, the final outcome. Different studies describe a reduction in cell

proliferation or differentiation, decreased ECM production, or increased ROS production

due to ageing or chronic diseases including diabetes and cardiovascular diseases [175,

276-278]. In case of very high levels, these ROS products might influence neo-tissue

generation or remodeling, as excessive levels of ROS products are toxic for cells [279]. In

addition, it is described that the ratio between endothelial progenitor cells and smooth

muscle progenitor cells is disturbed in diabetic patients, which might result in reduced

vascular repair capacity [280]. However, important to mention here is that although these

changes in cellular behavior are found in elderly people or patients, we demonstrated in

chapters 2 and 3 that our human cells, which were obtained from elderly cardiovascular

patients showed the potential to produce tissues with ECM values that were comparable

to native tissues. This suggests that even though cell behavior might change with age or

underlying diseases, they probably still have sufficient capacity to create tissues that are

close to native in terms of ECM components. Also, in vivo remodeling in young and adult

sheep, of in vitro cultured cardiovascular replacements demonstrated that ECM amounts

in both groups were approaching native values [60, 67, 73].

Although animal studies are mainly performed on healthy animal models, research using

diseased models would be of added value to investigate how this would affect tissue

remodeling, immune responses and scaffold resorption. The development of in vitro

Chapter 6

104

disease model systems for an initial screening and validation is even desired to circumvent

large animal trials using diseased animal models.

6.3.4 Future perspectives

The use of innovative 3D-printers in the field of tissue engineering is recently introduced

with several types of 3D-printers being developed and tested in order to generate

cardiovascular tissue replacements [281, 282]. The use of such a device could result in the

development of in vitro tissue models with more native-like geometries for research and

training purposes and in reproducible products. However, most research focused on

printing materials containing living cells [281, 283]. For the in situ approach, the polymers

of choice are not compatible yet with the 3D printers that are available today.

Furthermore, the time needed to produce e.g. a heart valve is still longer compared to the

electrospinning approach. In addition, when the polymers of choice are compatible with

3D printers, further research is needed to investigate whether these 3D printed products

can meet all requirements to be used as an in situ cardiovascular substitute e.g. fiber

diameter between 5-10 µm and porosity >80%.

Personalized treatments, especially in patients with many comorbidities might promote a

beneficial outcome. This could include administering of granulocyte colony stimulating

factors (G-CSF) to mobilize hematopoietic stem cells from the bone marrow to the blood

[284], increase levels of vitamins, cytokines or enzymes which are known to promote

tissue formation and remodeling [285-287], or decrease inflammatory reactions and high

ROS levels by administering antioxidant drugs or medical gasses [288, 289].

6.4 Conclusion

The work of this thesis provided insight into the interplay between scaffold degradation

rate and tissue formation, in terms of amount and composition, and tissue remodeling.

The effect of slow- and fast-degrading scaffold materials on tissue formation and

remodeling was investigated with the use of in vitro models. In situ tissue engineering is

an attractive approach in terms of off-the-shelf availability, costs, reproducibility, and

regulatory issues in comparison with other tissue engineering approaches that are

investigated. To maintain mechanical integrity of the bare scaffold after implantation until

the tissue is sufficiently mature to take over this role, good mechanical properties and

resorption characteristics of the chosen material are of high importance. With the know-

how gathered in this thesis we are able to take a step forward towards the appropriate

scaffold material for cardiovascular in situ tissue engineering purposes. First, we

demonstrated that tissues grown on fast-degrading scaffold materials resulted in

compaction of the constructs, due to traction forces exerted by the contractile cells in

combination with fast loss of mechanical integrity of the scaffolds, while tissues grown on

General discussion

105

6

slow-degrading materials showed organized structures and maintained their 3D

configuration. This shows the importance of using a slow-degrading scaffold material for

in situ replacement of cardiovascular tissues, as sufficient time is needed for tissue to fully

develop and mature prior to scaffold resorption by the body, in order to prevent

compaction or retraction. In addition, in this thesis in vitro models were used to unravel

degradation characteristics of different (supramolecular) materials. Scaffold degradation

rate was correlated to macrophage phenotype, with increased visual degradation by the

inflammatory-type. Electrospun PCL scaffold with a fiber diameter of 10 µm favored

macrophage polarization to the regenerative phenotype, which is suggested to improve

tissue outcome. The slow-resorbing supramolecular materials are promising for in situ

tissue engineering of cardiovascular tissues as they exhibit strong and elastic mechanical

properties, desired for the replacement of cardiovascular tissues. Lastly, as desired

resorption properties will vary between various applications, it is of importance that

resorption is controlled and that the scaffold properties can be tuned. Work of this thesis

demonstrated that by combining polymer backbones modified with supramolecular

motifs we are able to tune the scaffold properties and therewith are able to create

implants with the desired requirements, which can be further tested in pre-clinical

studies.

106

107

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127

Nederlandse samenvatting Wanneer een bloedvat of hartklep niet meer goed functioneert is het vaak noodzakelijk

om deze te vervangen door een prothese. Het grote nadeel is dat de beschikbare

prothesen geen levend weefsel bevatten en daardoor niet in staat zijn om zich aan te

passen aan veranderingen in hun omgeving. Hierdoor gaan ze niet levenslang mee of

dienen er levenslang antistollingsmiddelen te worden gebruikt. Om zieke weefsels te

vervangen met levende prothesen zijn er verschillende methodes ontwikkeld volgens het

principe van tissue engineering (TE). Een veelbelovende methode binnen dit principe is de

in-situ TE methode. Hierbij wordt een biologisch afbreekbare mal, ook wel scaffold

genoemd, in de vorm van een bloedvat of hartklep in het lichaam geplaatst om vervolgens

door het lichaam te worden getransformeerd in een levende prothese. Tijdens de vorming

van weefsel in en rondom de scaffold, breekt de scaffold af en blijft er gezond, levend

weefsel over dat gemaakt is door het lichaam van de patiënt. Het is van groot belang om

het juiste materiaal te gebruiken zodat de scaffold na implantatie mechanisch stabiel is

en pas afbreekt wanneer er voldoende weefsel is opgebouwd om deze functie over te

nemen. Dit proefschrift beschrijft een zoektocht naar het optimale scaffold materiaal

waarbij de interactie tussen de afbraaksnelheid van de scaffold en de eigenschappen van

het gevormde weefsel zijn onderzocht. Daarnaast zijn de afbraak eigenschappen van

scaffolds gemaakt van supramoleculaire materialen bestudeerd.

Wanneer de balans tussen het afbreken van de scaffold en de weefselvorming verstoord

is kan de mechanische stabiliteit niet gegarandeerd worden. Dit heeft in het verleden

geresulteerd in het samentrekken (compactie) en krimpen (retractie) van de vliesjes van

in-vitro gekweekte hartkleppen, met lekkende kleppen tot gevolg. We hebben bestudeerd

of dit voorkomen kon worden door het gebruik van een langzaam-afbrekend scaffold in

plaats van een snel-afbrekend scaffold en/of het gebruik van een lager cel passage

nummer om de weefselontwikkeling te bevorderen. Compactie en retractie

verminderden wanneer er gebruik werd gemaakt van langzaam afbrekend scaffold.

Verder bleken de weefselvorming, stijfheid en sterkte toe te nemen met afnemende cel

passage nummers, wat daarentegen geen invloed had op de compactie en retractie.

Verder is het effect van afbraaksnelheid van scaffolds op de hoeveelheid en samenstelling

van het gevormde weefsel onderzocht, evenals de mechanische stabiliteit en de

verhouding tussen gevormd weefsel en overgebleven scaffold. Hierbij werd wederom

gebruik gemaakt van scaffolds met langzame en snelle afbraak. De hypothese hierbij was

dat de cellen in het snel-afbrekende scaffold meer weefsel zouden aanmaken, in

verhouding tot cellen in het langzaam-afbrekende scaffold, om te compenseren voor het

snelle verlies van de mechanische stabiliteit. In beide materialen nam de hoeveelheid

weefsel toe tijdens de kweektijd met de grootste toename in het snel-afbrekende

Nederlandse samenvatting

128

materiaal, echter alleen tijdens de eerste 2 weken. Na 6 weken kweken bleken de

langzaam-afbrekende materialen de hoogste hoeveelheden weefsel te bevatten. Om een

eerlijk vergelijk te kunnen maken tussen de weefselsamenstelling van TE weefsels en

natieve hartkleppen is er ook een methode beschreven om te corrigeren voor het gewicht

van het resterende scaffold. Na het toepassen van deze correctie bleek dat de weefsel

componenten in de TE weefsels vergelijkbaar waren met die van natieve pulmonaal

hartkleppen. Een uitzondering hierop vormde de collageen crosslinks. Ondanks een

toename van collageen crosslinks in de TE weefsels met de tijd was deze hoeveelheid lager

dan die in natieve weefsels.

In het lichaam kunnen scaffolds afgebroken worden via de oxidatieve en/of de

hydrolytische route. Om beide routes apart en op een versnelde manier te onderzoeken

zijn in-vitro testen ontworpen. De afbraak eigenschappen van scaffolds gemaakt van

verschillende veelbelovende supramoleculaire materialen zijn onderzocht en vergeleken

met scaffolds gemaakt van het conventionele materiaal polycaprolactone (PCL). Scaffolds

reageerden anders op beide afbraak routes, afhankelijk van de morfologische en

chemische compositie van het materiaal. PCL is gevoeliger voor hydrolyse vergeleken met

supramoleculaire materialen, terwijl het tegenovergestelde werd gezien tijdens de

oxidatieve afbraak route. Daarnaast bleek het mogelijk om de afbraaksnelheden te

beïnvloeden. Op basis hiervan kunnen veelbelovende materialen in de toekomst worden

gescreend, en het optimale materiaal worden geselecteerd voor preklinische studies.

Na implantatie van een scaffold komt er een immuunreactie op gang waarbij er o.a.

macrofagen de scaffold binnen dringen. Deze macrofagen spelen een belangrijke rol in

zowel het afbreken van het scaffold als in de opbouw van weefsel. Er zijn verschillende

typen macrofagen, waaronder een ontstekings-bevorderend en een helingsbevorderend

type. Het is echter nog onduidelijk wat de bijdrage van deze macrofaagtypen op de

afbraak van materialen is. In dit proefschrift is onderzocht of er een verband is tussen

macrofaagtype en afbraak van scaffold. Dit laat zien dat het macrofaagtype van invloed is

op de afbraak van scaffold, waarbij de ontstekingsbevorderende macrofagen een

versnelde afbraak lieten zien. Daarnaast bleek dat het PCL materiaal de macrofagen

richting het helingsbevorderende type stuurde, wat gunstig is voor gebruik in in-situ TE.

Samengevat is de keus van het scaffold materiaal belangrijk om de mechanische stabiliteit

te behouden. De resultaten van dit proefschrift benadrukken dat een langzaam afbrekend

scaffold gewenst is, omdat de mechanische stabiliteit dan langer behouden blijft, wat erg

belangrijk is voor in-situ TE doeleinden. Daarnaast bleek dat weefsel beter georganiseerd

was en tot minder compactie leidde wanneer dit gekweekt werd op langzaam afbreekbare

scaffolds. Resultaten in dit proefschrift laten tevens zien dat de afbraaksnelheid van

scaffolds te beïnvloeden is, wat van belang is omdat verschillende afbraak snelheden

gewenst zijn bij verschillende toepassingen.

129

Dankwoord Zo, het zit er alweer zo goed als op. De experimenten op het lab zijn uitgevoerd, de resultaten

verwerkt, het proefschrift geschreven, en de papers gepubliceerd, onderweg of in de maak. Ik ben

nu echt (bijna) ‘afgestudeerd’ en heb na meer dan 4 jaar (bijna) ‘promotie’ gemaakt. Dit zou ik

nooit alleen voor elkaar hebben gekregen, maar gelukkig waren er afgelopen jaren vele

hulptroepen die mij, zonder het soms zelf te weten, super veel hebben geholpen. Daarom wil ik

al deze personen graag bedanken.

Martijn, wat een mooie kans heb je me geboden dat ik binnen Xeltis een promotie onderzoek

mocht uitvoeren. Bedankt voor de hulp afgelopen jaren, de nuchtere adviezen en de vrijheid die

ik altijd heb gekregen om mijn promotie tussen de werkzaamheden van Xeltis door te kunnen

plannen. Het was super leuk en leerzaam om tijdens mijn promotie ook mee te kunnen maken

hoe Xeltis de afgelopen jaren is gegroeid. Ik ben benieuwd hoe we verder ontwikkelen.

Frank, bedankt dat je de afgelopen jaren mijn 1e promotor wilde zijn. Bedankt ook voor je slimme

opmerkingen tijdens de besprekingen waardoor ik weer vooruit kon met het onderzoek. We

hebben beiden ondervonden dat er soms verschillende belangen binnen een bedrijf en

universiteit spelen. Ondanks dat dit niet altijd makkelijk was heb ik er ook veel van geleerd. Ik ben

blij en dankbaar dat we uiteindelijk tot dit proefschrift zijn gekomen. Heel veel succes in je nieuwe

functie als Rector Magnificus. Carlijn, met name het laatste deel van mijn promotie had ik de luxe

om gebruik te mogen maken van jouw waardevolle kennis en begeleiding als 2e promotor. Dit heb

ik erg gewaardeerd en hier heb ik ook veel aan gehad. Bedankt voor de input en het lezen van alle

stukken.

Anita, van jouw structurele aanpak tijdens het analyseren van de bergen aan data en manier van

schrijven heb ik echt super veel geleerd. Met de smiley’s die ik vaak tegenkwam op de

gecorrigeerde versies van mijn papers of hoofdstukken werd het verbeteren hiervan een stuk

leuker . Ik vond het erg fijn dat je mijn co-promotor wilde zijn. Bedankt!.Ik wens je heel veel

succes met je nieuwe uitdaging, maar voorlopig ben je ook nog op de TU/e te vinden. Gezellig!

Peter Hilbers, Patricia Dankers, Pamela Habibovic, Christian Ottmann, en Jolanda Kluijn hartelijk

bedankt voor het deelnemen in mijn commissie.

Mirjam, zeker ook mede dankzij jou heb ik de kans gekregen om een promotie onderzoek te doen.

Met dit proefschrift als resultaat. Bedankt voor al je hulp en adviezen tijdens het grootste deel van

mijn promotie. Fijn om te horen dat je het in je nieuwe functie naar je zin hebt.

Dear Xeltis colleagues in Eindhoven and Zurich, thank you very much for the great work

atmosphere. There is always some laughter, even in times of deadlines and very busy working

days. I think it is really fun working with all of you. Special thanks to Anita and Leonie for all the

SEM and tensile tests you performed. Tom, thanks for all the questions I could ask related to

mechanical tests. Roel and Marc, thank you for all the scaffolds you prepared when I needed

material for experiments. Anand and Ioannis, thank you for reading some of the chapters. Within

Xeltis everybody works very hard, but we also often get the opportunity to play hard, which I

greatly appreciate. The ‘heerlijk avondjes’ where the crazier you dress up the better, the dinners,

driving a tank ourselves, waterskiing, shooting with a shotgun, and building our own little igloo

Dankwoord

130

town, in which we even could stay overnight were all great moments. I really enjoy all these

activities and moments we spent together.

Ook heb ik met veel plezier een deel van mijn promotie doorgebracht op de TU/e. Voor lab zaken

kon ik terecht bij Moniek, voor praktische zaken stond Yvon altijd klaar. Zo waren er nog velen

meer met wie ik fijn heb samengewerkt, of nog steeds doe in het Ivalve project. Bedankt! Dear ex-

4.12-roomies, it was really good fun with you during working hours, the coffee breaks, and during

the 4.12 room outings. Thanks a lot for the good times. Henk, Serge en Tonny, bedankt voor de

fijne samenwerking en input voor het vierde hoofdstuk van dit proefschrift.

Gelukkig is er nog meer dan alleen werk. Ik geniet dan ook altijd goed van de tijd die ik samen met

familie en vrienden door kan brengen. Dit zorgt voor de leuke ontspanning naast het werk.

Kristie, Anke, Carolien, Janneke en Willeke, vriendinnen uit de “Hei”. We kennen elkaar bijna

allemaal al vanaf groep 2 toen ik bij jullie in de klas kwam. Sindsdien hebben we samen echt veel

meegemaakt. De jaarlijkse weekendjes weg, de avondjes stappen, high tea’s, etentjes, borrels enz.

zijn altijd top. Behalve dat het vaak een leuk kippenhok is met jullie, is het ook super fijn om alles

met jullie te kunnen delen. Bedankt meiden! Met Twan, Geoffrey, Ruud U., Mark, Maarten, Karin

en Giel, Wendy en Ruud H. heb ik ondertussen ook al menig uurtjes op elk moment van de

dag/nacht mogen slijten. Bedankt voor de slappe klets, gezelligheid, ontspanning en interesse

tijdens al deze uurtjes.

Hetty, de avondjes in Nijmegen waren altijd een groot feest! Net zoals de borrels die we doen in

Valkenswaard als je vakantie kan nemen op Yale. Bedankt voor alle tips en trucs die je me hebt

gegeven tijdens de afronding van mijn promotie. Anneke, Martijn & Judith, Kirsten & Wil, Rick &

Pam, buren in de Slechtvalk, bedankt voor de leuke feesten en etentjes. Bart, Sonja, Marc, Joyce,

Tanja en Alain, bedankt voor de gezelligheid tijdens carnaval, etentjes, tonproaters avonden,

borrels en weekendjes weg. Ik vind het altijd super leuk om dingen met jullie te ondernemen.

Dames van korfbalclub de Stormvogels, allemaal bedankt voor de leuke, sportieve tijd afgelopen

jaren. Niks is zo ontspannend als jezelf een paar keer per week, na een dagje werken, lekker leeg

te rennen op het veld. Het liefst gevolgd door gezellig wat nakeuvelen in de kantine waar we er

vaak nog meer calorieën bij eten/ drinken dan we er van tevoren afgesport hadden. Heerlijk!

En wat doe je als je na het werk wat wilt ontspannen maar het korfbal seizoen is afgelopen? Een

bloemencorso wagen bouwen natuurlijk! Dankzij alle “hazen” van Buurtschap Hazestraat, van

jong tot oud, heb ik de afgelopen paar jaar met veel plezier geholpen bij de bar, de figuratie en de

wagen. Wat is het fijn om deel uit te mogen maken van zo’n mooi clubke, bedankt allemaal!

Pap, Mam, Debbie, Wouter, Joost, Hans, Lisette, Piet, Marianne, Debbie, Rob en Naomi, bedankt

voor alle steun, de fijne tijd die we samen hebben en de interesse die jullie altijd in mij getoond

hebben. Jullie zijn een hele fijne (schoon)familie waar ik erg blij mee ben!

Robin, het was echt super fijn om de afgelopen jaren alle dingen die goed gingen, maar ook de

dingen die niet volgens plan verliepen samen met jou te kunnen delen. Bedankt dat je altijd achter

me staat, en voor al je steun en vertrouwen. Het is fijn om met jou samen te wonen, te genieten

van het leven, en vooral veel te lachen. Ik ben kei gelukkig met jou! Op naar een mooie toekomst!

Marieke, april 2015

131

Curriculum vitae

Marieke Brugmans is geboren op 25 november 1984 in Veghel. In 2003 behaalde zij haar

VWO diploma aan het Zwijsen College in Veghel. Aansluitend studeerde zij Biomedische

Wetenschappen aan de Radboud Universiteit te Nijmegen. Als onderdeel van haar studie

heeft ze 5 maanden stage gelopen bij het Centraal Hematologisch Laboratorium, op de

afdeling stamcel transplantatie en immuun therapie aan het UMC St. Radboud ziekenhuis

te Nijmegen. Hier heeft ze een bijdrage geleverd aan de ontwikkeling van een micro-array

voor screenings doeleinden. Vervolgens heeft ze 7 maanden stage gelopen bij Organon in

Oss, waar ze heeft geholpen met het opzetten van een modelsysteem voor het

voorspellen van behandel effecten op genezing van bot fracturen. Hierna heeft ze 5

maanden onderzoek verricht bij het Nederlands Kanker Instituut in Amsterdam waarbij

de focus lag op de transfectie en retrovirale transductie van cellijnen om meer inzicht te

krijgen in de rol van CXCL10 op T-cellen. In 2008 heeft zij haar afstudeeronderzoek van 9

maanden gedaan bij het Diamantia Instituut voor kanker, immunologie en metabole

ziekten in Brisbane, Australië. Hier heeft ze onderzoek verricht naar het identificeren van

genen die een effect kunnen hebben op de anabole controle van botgroei door middel

van genotypering. In 2009 behaalde zij haar master diploma in de richting humane

pathobiologie. Vervolgens is ze begonnen als research technician bij Xeltis B.V. en hier een

jaar later gestart met haar promotieonderzoek in samenwerking met de vakgroep Soft

Tissue Biomechanics & Engineering aan de faculteit Biomedische Technologie van de

Technische Universiteit Eindhoven, resulterend in dit proefschrift. Op het moment is ze

werkzaam als research engineer bij Xeltis B.V. in Eindhoven.

133

List of publications

Peer-reviewed publications

M Brugmans, A Driessen-Mol, M Rubbens, M Cox, F Baaijens, Poly-ε-caprolactone scaffold and reduced in vitro cell culture: beneficial effect on compaction and improved valvular tissue formation, J Tissue Eng Regen Med. 2013 (in press)

V Peperzak, E Veraar, Y Xiao, N Babala, K Thiadens, M Brugmans, J Borst, CD8+ T cells produce the chemokine CXCL10 in response to CD27/CD70 costimulation to promote generation of the CD8+ effector T cell pool, J Immunol. 2013 Sep;191(6):3025-36.

N de Jonge, J Foolen, M Brugmans, S Söntjens, F Baaijens, C Bouten, Degree of scaffold degradation influences collagen (re)orientation in engineered tissues, Tissue Eng Part A. 2014 Jun;20(11-12):1747-57

M Brugmans, R Soekhradj-Soechit, D van Geemen, M Cox, C Bouten, F Baaijens, A Driessen-Mol, Superior tissue evolution in slow-degrading scaffolds for valvular tissue engineering (submitted)

M Brugmans, S Sӧntjens, M Cox, A Nandakumar, A Bosman, T Mes, H Janssen, C Bouten, F Baaijens, A Driessen-Mol, Hydrolytic and oxidative degradation of electrospun supramolecular biomaterials: In vitro degradation pathways. (submitted)

M Brugmans, M Cox, C Bouten, F Baaijens, A Driessen-Mol, Advanced electrospun scaffold degradation by inflammatory macrophages in comparison with healing macrophages (in preparation)

C Gregson, L Wheeler, S Hardcastle, J Pointon, K Addison, M Brugmans, G Clark, K Ward, M Paggiosi, J Turton, M Stone, J Thomas, R Agarwal, K Poole, E McCloskey, E Williams, A Bullock, G Smith, M Brown, J Tobias, E Duncan, Predictions in LRP5 protein structure explain variation in the clinical severity of LRP5 High Bone Mass. (submitted)

Peer-reviewed proceedings

M. Brugmans, A. Driessen-Mol, M. Cox, M. Rubbens, and F. Baaijens, PCL scaffolds and reduced in vitro cell expansion to improve engineered valvular tissue formation. 5th biennial conference on heart valve biology and tissue engineering. Oral presentation, Myconos, Greece, 2012

M. Brugmans, S. Sӧntjens, M. Cox, A. Nandakumar, A. Bosman, T. Mes, H. Janssen, C. Bouten, F. Baaijens and A. Driessen-Mol, Hydrolytic and Oxidative Degradation of Electrospun Supramolecular Biomaterials: In Vitro Degradation Pathways. 26th annual conference of the European Society of Biomaterials. Oral presentation, Liverpool, England, 2014

http://www.ncbi.nlm.nih.gov/pubmed?term=Driessen-Mol%20A%5BAuthor%5D&cauthor=true&cauthor_uid=23677869

http://www.ncbi.nlm.nih.gov/pubmed?term=Rubbens%20MP%5BAuthor%5D&cauthor=true&cauthor_uid=23677869

http://www.ncbi.nlm.nih.gov/pubmed?term=Cox%20MA%5BAuthor%5D&cauthor=true&cauthor_uid=23677869

http://www.ncbi.nlm.nih.gov/pubmed?term=Baaijens%20FP%5BAuthor%5D&cauthor=true&cauthor_uid=23677869

http://www.ncbi.nlm.nih.gov/pubmed/23677869

http://www.ncbi.nlm.nih.gov/pubmed?term=Peperzak%20V%5BAuthor%5D&cauthor=true&cauthor_uid=23940275

http://www.ncbi.nlm.nih.gov/pubmed?term=Veraar%20EA%5BAuthor%5D&cauthor=true&cauthor_uid=23940275

http://www.ncbi.nlm.nih.gov/pubmed?term=Xiao%20Y%5BAuthor%5D&cauthor=true&cauthor_uid=23940275

http://www.ncbi.nlm.nih.gov/pubmed?term=Babala%20N%5BAuthor%5D&cauthor=true&cauthor_uid=23940275

http://www.ncbi.nlm.nih.gov/pubmed?term=Thiadens%20K%5BAuthor%5D&cauthor=true&cauthor_uid=23940275

http://www.ncbi.nlm.nih.gov/pubmed?term=Brugmans%20M%5BAuthor%5D&cauthor=true&cauthor_uid=23940275

http://www.ncbi.nlm.nih.gov/pubmed?term=Borst%20J%5BAuthor%5D&cauthor=true&cauthor_uid=23940275


http://www.ncbi.nlm.nih.gov/pubmed?term=de%20Jonge%20N%5BAuthor%5D&cauthor=true&cauthor_uid=24372199

http://www.ncbi.nlm.nih.gov/pubmed?term=Foolen%20J%5BAuthor%5D&cauthor=true&cauthor_uid=24372199

http://www.ncbi.nlm.nih.gov/pubmed?term=Brugmans%20M%5BAuthor%5D&cauthor=true&cauthor_uid=24372199

http://www.ncbi.nlm.nih.gov/pubmed?term=S%C3%B6ntjens%20S%5BAuthor%5D&cauthor=true&cauthor_uid=24372199

http://www.ncbi.nlm.nih.gov/pubmed?term=Baaijens%20F%5BAuthor%5D&cauthor=true&cauthor_uid=24372199

http://www.ncbi.nlm.nih.gov/pubmed?term=Bouten%20C%5BAuthor%5D&cauthor=true&cauthor_uid=24372199



The interplay between biomaterial degradation and tissue … › ws › files › 24302469 › 20160610_Brugmans.pdf · 6.2 Towards the most promising tissue engineering approach

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