tissue engineering in congenital heart...

Challenges in Regenerative Medicine Vol.1 No.1 August 2014

ISSN: xxxx-xxxx (Print) ISSN: xxxx-xxxx (Online) http://www.researchpub.org/journal/iphf/iphf.html

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Abstract— Congenital heart disease is one of the most

prevalent birth defects. Approximately 1% of babies are born

with a structural heart defect. Components of the

cardiovascular system that may be affected by congenital heart

disease include the cardiac wall, septum, heart valves and blood

vessels. Many of these cardiac anomalies require surgical

correction; however current methods have some limitations.

Pediatric patients often require multiple operations as they

grow to replace their grafts, which may be either failing or

becoming too small in size. Additionally, current materials for

organ reconstruction can still not fully integrate with the host

tissue and lack regenerative properties. Tissue engineering

potentially offers a solution to the pressing problem of tissue

damage and shortage of transplant organs. This review article

provides an overview of congenital heart defects and outlines

the unique needs of children as a patient population. Advances

in tissue engineering of cardiac patches, heart valves and blood

vessels relevant to congenital heart disease are described.

Furthermore, this report outlines the challenges in the field that

need to be overcome before tissue engineered myocardium,

septum, valves and vessels can become clinical reality for the

treatment of pediatric cardiac patients.

Keywords — Biomaterial, cardiac defect, congenital heart

disease, regenerative medicine, tissue engineering

1. INTRODUCTION

Congenital heart defects (CHDs) affect approximately 1 in

100 children born in the United States, which equates to nearly

40,000 births per year [1, 2]. These statistics are a conservative

estimate, and can vary depending on the classification and

detection of the defect [1]. As a result, CHDs are the leading

cause of infant illness and death [3]. From 1999 to 2006, the

US census reported nearly 41,500 CHD-related deaths with

almost 28,000 deaths resulting directly from CHDs [4]. While

the exact mechanisms that cause each CHD are unclear, it is

estimated that at least 15% are related to certain genetic

disorders [5-7]. The type and severity of defects are vast and

often occur as a combination of defects [8]. The spectrum of

CHDs is very broad, with some minor defects not requiring any

form of therapy, to very complex cases, where time-sensitive

1 Department of Pediatrics, Baylor College of Medicine, One Baylor Plaza,

Houston, TX, 77030, USA, 2 Department of Bioengineering, Rice University,

PO Box 1892, MS-142, Houston, TX 77251-1892, USA 3 Congenital Heart

Surgery Services, Texas Children’s Hospital, Houston, TX, USA

*Correspondence to Jeffrey G. Jacot (e-mail: [email protected]).

surgical interventions are needed to sustain the life of the child.

The cost of healthcare for CHDs is massive, with nearly $1.4

billion reported in the United States in 2004 [9]. This is more

than half of the entire hospital costs for birth defects reported

that year. Therefore early detection and strategies to repair

CHDs and prevent long-term therapy and multiple operations

are important areas of research that would have significant

return.

In recent years, tissue engineering has emerged as a major

field of research for the regeneration and restoration of injured

or defective native tissue. Where some regenerative success in

years prior has come from autografts and homografts [10, 11]

restoring structure similar to native tissue, the majority of

approaches lack the restoration of native function. In the case

of CHDs, especially those involving cardiac wall repair, the

ability to contract synchronously is essential for proper blood

circulation. Tissue engineering is an appealing alternative that

can overcome the shortfalls of grafts in terms of availability and

rejection if used in conjunction with autologous cells. Tissue

engineered constructs for CHDs would not only provide

structural support, but also restore native function. Therefore

classical tissue engineering strategies [12, 13] to accomplish

this goal are typically composed of the following key elements:

a scaffold for mechanical support, mechanical stimuli to

provide conditioning, biochemical stimuli to promote cellular

signaling and function, and cells to assist in new tissue

formation as well as signaling. Ideally, the construct would

also be biodegradable with the degradation rate in tune with the

new tissue ingrowth.

The developing cardiopulmonary system is unique and

dynamic. This system is essential to sustaining life, and does so

by means of its structural and functional complexity. Cardiac

tissue is unique in itself, where cardiac muscle is distinct from

skeletal and smooth muscle. Components of the cardiovascular

system are regionally and functionally different, containing

different cells and extracellular matrix molecules.

Understanding the complexity of this system is the first step

towards designing regenerative techniques. In the case of

pediatric patients, CHDs pose further unique issues compared

to adult chronic cardiovascular diseases, such as

atherosclerosis, and therefore one must also consider the

distinctive needs of children as a patient population.

Many congenital heart defects are detected during prenatal

testing [14], or diagnosed during the first days of life. The most

severe lesions may require treatment very soon after birth. The

timing of intervention is a critical step, both in terms of surgical

decision-making, but also because of the availability of tissue

grafts. Sometimes a procedure cannot be undertaken until a

donor tissue of the right size is found, or because a surgeon

elects to postpone treatment until the patient grows and can

receive a larger implant so that multiple repeat surgical

procedures can be avoided. Tissue engineering could

Tissue Engineering in Congenital Heart Disease

Leda Klouda1, PhD, Christopher Tsao

2, MS, and Jeffrey G. Jacot*

2, 3, PhD

http://www.researchpub.org/journal/iphf/iphf.html



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circumvent these concerns by providing a made-to-order

solution that can grow with the child. In the case that cells will

be included in the tissue engineered construct, prenatally

harvested cells could be expanded and be ready for

implantation quickly after the baby is born. Also, considering

the young age of patients receiving a graft, it is important to

stress the unique need for the implant to remodel and grow

during the child’s stages of development. A child undergoes

incremental growth phases, and the increase in size of most

organs follows that of the general body size [15]. Ideally, a

construct should be fully integrated with the host and be

durable and functional for life. Studies are presently underway

to determine the life span of some solutions employed for the

treatment of congenital cardiovascular disease, and in many

cases, reoperations and replacements are necessary. For

example, existing surgical prostheses for heart valve disease

have a limited functional life, which may be adequate for a

middle-aged person, but often prove problematic for young

recipients. Moreover, differences in the immune system of

infants, children and adults must be carefully evaluated

[16-18]. Previous studies have suggested that the performance

of bioprostheses is linked to the immune response which can

vary between age groups, particularly infants [19-21]. The

latter age category is important due to the large volume of

congenital cardiac surgeries performed in early infancy.

Therefore the use of immunosuppressants in pediatric patients

needing bioprostheses becomes a delicate balance of implant

rejection, side effects of long term exposure, and risk of

infection. When designing treatment strategies, one must

consider the detrimental effects of certain compounds or

medications on the developing systems of young children, and

also their active lifestyles, which make anticoagulant drugs for

example often dangerous. Overall, these are some of the factors

that may differentiate management of pediatric versus adult

patients, and have potential implications in tissue engineering.

This review provides an overview of the most common

congenital heart defects, as well as highlight current tissue

engineering research strategies used to correct these defects.

From this literature examination, current challenges in the field

are elucidated and therefore give insight into future research

directions for more successful treatment strategies.

2. OVERVIEW OF CONGENITAL HEART DEFECTS

Based on clinical studies, CHD is believed to affect around

1% of live births [1]. Ultrasound procedures allow for

approximately 39% of major CHDs, such as tetralogy of Fallot

or hypoplastic left heart syndrome, to be diagnosed in utero

during the second trimester [14], while the diagnosis of acute

CHDs, such as small atrial septal defects (ASDs) or ventricular

septal defects (VSDs), may not occur until birth or later.

Typically a trivial cardiac lesion that has eluded detection will

naturally close during infancy [22]. Also, many CHDs do not

manifest themselves until there is mixing of oxygenated and

deoxygenated blood after birth. This results from incomplete

closure of fetal circulatory vessels, such as the ductus

arteriosus. Therefore the actual reported percentage of births

with CHDs is likely to be slightly higher than reported [1].

The heart is a complex organ and myriad diseases can

result from lesions or malfunctions. As suggested by the

statistical occurrence of CHDs, the degree of severity in CHDs

is very diverse. There are a number of risk factors that can lead

to CHD, including but not limited to, maternal diabetes and

maternal lithium, phenytoin and alcohol use. In cases of

maternal diabetes, the risk of having a newborn with a

structural heart defect increases by 30% [23]. Family history

of CHD can also play a role, about 1-4% of babies born to

parents with CHDs are affected [1]. The exact causes for the

majority of CHDs are unknown, but are proposed to be linked

to causative genes in developing fetuses [7]. From the

perspective of tissue engineering for CHDs, corrective action

attempts to treat different lesions associated with a given

congenital heart disease. A brief overview of clinically

relevant lesion specific CHDs will help to better understand the

engineering challenges in creating a viable construct. Table 1

summarizes some specific CHDs with associated causes and

current treatment options.

TABLE OF ABBREVIATIONS AFSC amniotic fluid derived stem cells

ASD atrial septal defect

CHD congenital heart defect/disease

ECM extracellular matrix

EPC endothelial progenitor cells

GAG glycosaminoglycan

HA hyaluronic acid

IGF insulin growth factor

MSC mesenchymal stromal cells

P4HB poly(4-hydroxybutyrate)

ePTFE expanded polytetrafluoroethylene

PA pulmonary artery

PCL polycaprolactone

PDA patent ductus arteriosus

PEG poly (ethylene glycol)

PEUU polyester urethane urea

PGA poly(glycolic acid)

PGS poly(glycerol sebacate)

PLGA poly(lactide-co-glycolide)

PLLA poly(L-lactic acid)

PU polyurethane

RVOT right ventricular outflow tract

SIS small intestinal submucosa

TEHV tissue engineered heart valve

VEGF vascular endothelial growth factor

VIC valvular interstitial cell

VSD ventricular septal defect




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Table 1: Congenital Heart Defects

Defect Estimated

% of Total

Congenital

Heart

Defects in

the US [1,

2]

Causes Diagnosis Treatment

Options

Ventricular

septal defects

20% Hole in ventricular

septum

Cardiac auscultation,

chest x-ray,

echocardiogram

Occlusion devices

[24-26],

patch closure materials

[27]

Tetralogy of

Fallot

10% 1. Large

ventricular defect 2.

Severe right

ventricular outflow tract

obstruction 3.

Overriding of the aorta 4.

Right

ventricular hypertrophy

Echocardiogra

m, chest x-ray, pulse

oximetry, MRI

Right

ventricular outflow tract

reconstructio

n [28] VSD closure

device [29,

30]

Coarctation

of the aorta

10% Mild-complete

obstruction of

aorta

Magnetic

resonance

angiography, echocardiogra

m

Patch for

aortoplasty

[31]

Patent

ductus

arteriosus

10% Ductus arteriosus

remains open

after birth

Chest x-ray, echocardiogra

m

PDA occlusion

device [32],

occlusion spring coil

[33]

Atrial septal

defect

5% Hole in atrial

septum

Echocardiogra

m, cardiac

auscultation, chest x-ray,

Occlusion

devices [34, 35], patch

closure

materials [36, 37]

Transpositio

n of the

great vessels

5% Abnormal

spatial

arrangement of great

vessels

Echocardiogra

m, X-ray

Arterial

switch

operation [38], septal

occlusion

device

Aortic valve

stenosis

5% Narrowing across aortic

valve

Echocardiogram

Surgical or balloon

valvuloplast

y [39], bioprosthese

s [40]

Ebstein

anomaly

1% Abnormal

tricuspid valve and

atrialization of

ventricle

Echocardiogra

m, ECG, chest x-ray

Valve

reconstruction,

Bioprostheti

c or mechanical

valve,

surgical or transcatheter

placement

[41, 42]

Anomalies

of the

1% Mild-complete obstruction of

Pulse oximetry,

Venous connection

pulmonary

veins

pulmonary

artery

Cardiac

catheterization

reconstructio

n, patch

closure materials

[43, 44]

Tricuspid

atresia

1% Absence of

tricuspid valve

Echocardiogra

m, chest x-ray, ECG

Valve

reconstruction,

Bioprostheti

c or mechanical

valve,

surgical or transcatheter

placement

[42, 45]

Interrupted

aortic arch

1% Underdeveloped aorta

Echocardiogram, loss of

appetite, pale

Patch for aortoplasty

[31]

Hypoplastic

left heart

syndrome

1% Varying

degrees of underdevelope

d aorta, aortic

valve, left ventricle,

mitral valve,

and left atrium

Echocardiogra

m

Left heart

total repair, shunts [46],

aortoplasty

[31]

3. TISSUE ENGINEERING STRATEGIES FOR CARDIAC WALL

REPAIR IN CONGENITAL HEART DISEASE

In congenital heart disease, common abnormalities requiring

surgical intervention involve a lesion or hole in septal heart

tissue requiring closure. Small defects (<5mm) often do not

require surgical intervention and will spontaneously close [54].

However, when the hole is too large, surgeons will place

patches or occlusion materials in a patient to help bridge the

native heart tissue and allow for restored structure. In terms of

an ideal tissue engineering construct, restoring structural

support is only part of the challenge. A biomaterial may meet

the mechanical properties needed to withstand the fatigue of a

contracting heart, but if the material is inert it will not

completely restore native function. Vascularization is also a

limiting factor in current cardiac patches [55]. Unless the

cardiac patch is prevascularized or able to have neovascular

ingrowth, the patch will be diffusion limited and therefore

limited to the overall thickness [56]. Another requirement in

restoring native function is the synchronized beating of an

implanted cardiac patch with the native heart pace [57].

Following the restoration of natural function is the ability of

living tissue to grow and expand. This is particularly important

in pediatric patients, where their continued growth can lead to

implant failure, requiring reoperations [22]. This necessity to

grow and expand presents a unique engineering challenge

which is not as relevant in adult applications. A number of

CHDs require cardiac wall repair, the most common being

ventricular/atrial septal wall repairs using occlusion devices

termed “baffles”, which are defined as materials that direct the

flow of blood between desired chambers while sealing blood

from other chambers [58]. This is contrary to full-thickness

cardiac wall patches where the goal is to augment free wall

structures. Examples of full-thickness cardiac patches needed

during surgical intervention include: ventricular outflow patch




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for tetralogy of Fallot, complete left heart remodeling in

hypoplastic left heart syndrome, and right ventricular patch in

Ebstein anomaly. To date, no commercially available patch can

completely restore native structure and function after

implantation.

Naturally derived cardiac patches

A number of strategies involve the use of naturally derived

materials as a scaffold for a cardiac tissue engineering

construct. The major advantage of naturally derived materials

is that they are composed of molecules found in vivo and will

degrade into natural metabolic products. Also cellular response

can be favorable compared to other synthetic materials in that

adhesion molecules which enhance bioactivity are already

present within the natural material. In CHDs, many surgical

procedures have used autologous pericardium as patch material

[59]. The advantages to this technique are that the material is

immediately available, nonimmunogenic, and free of cost.

Pericardium has mechanical properties that are inferior to

native cardiac tissue and therefore in order to enhance the

mechanical properties, some surgical procedures involve

glutaraldehyde fixation of the material. With this method,

aldehyde groups will crosslink at the amine groups on the

lysine and hydroxylysine residues of pericardium collagen.

Pericardium is typically crosslinked for 15 to 30 minutes,

where the duration of crosslinking can affect the resulting

mechanical properties [60]. Glutaraldehyde is a toxic solution

and therefore adequate washing of the patch material is needed

prior to implantation [61]. Other crosslinking agents, such as

genipin [62] and acyl azide [63], have been tested as well. One

of the main drawbacks in the use of autologous pericardium as a

patch material is the inability of the material to expand or grow

in pediatric patients. Also, crosslinking procedures will lyse

any native cells, as well as create a chemically different

material resulting in possible calcification and fibrous

encapsulation [64].

The use of decellularized matrix as a scaffold is appealing

because the construct was originally functional tissue.

Depending on the decellularization process, key ECM

molecules can be preserved and therefore may help in

promoting new tissue remodeling. Currently there are three

decellularized patches approved for cardiovascular patch

applications; a pulmonary artery patch material

(MatrACELLTM , Virginia Beach, VA), a pericardial patch

(CryoPatch®, CryoLife Inc, Kennesaw, GA), and a

decellularized porcine intestinal submucosa (CorMatrix®,

Alpharetta, GA) [65]. CorMatrix® however is approved only

for use on artery, valve and pericardial tissue repair. Some

disadvantages associated with decellularized matrices are

immunogenicity, risk of disease transmission, and donor

availability [66]. Recent studies by Rajabi-Zeleti et al. [67]

attempt to renew the use of pericardium-derived patches using

an alternative approach to glutaraldehyde fixation. Their group

used decellularized pericardium that was enzymatically

digested and then reformed into pericardium gels. They

showed that after one month of subcutaneous implantation in

rats, the pericardium gel based scaffold had low immunological

response, enhanced angiogenesis, and cardiomyocyte

differentiation compared to control collagen and plain

decellularized pericardium. Crapo et al. [68] show that a small

intestinal submucosa (SIS) based gel seeded with neonatal rat

cardiac cells created tissue closer to physiological function

compared to that of cells seeded on Matrigel. Function was

measured in terms of contraction rate and normalized troponin

T expression, both higher in the SIS gel group. This difference

is attributed to the differing components of each gel. While

both gels contain a variety of different ECM components, the

main difference is that the SIS based gel contains high

concentrations of collagen type I and III, similar to that of

myocardium. At a minimum this study proves that myocardial

cell attachment and resulting function is complex and must be

considered carefully when choosing biomaterials for cardiac

tissue engineering.

Since native cardiac tissue extracellular matrix is

predominately composed of collagen, many studies aim to

manipulate collagen based scaffolds using growth factors

and/or altering properties such as alignment and microstructure

[69]. Collagen has good cellular attachment and proliferation

[70]. The main drawbacks to collagen based scaffolds in tissue

engineering applications are inferior mechanical and

degradation properties. A major requirement for cardiac tissue

engineering is the ability to withstand contractile forces of a

beating heart. Also as a patch material, collagen patches can be

difficult to suture due to their mechanical weakness. However,

since collagen is still the main component of cardiac ECM,

there is relevance to its use as a scaffold. Miyagi et al [71]

produced a collagen based patch that contained covalently

immobilized vascular endothelial growth factor (VEGF). This

patch was implanted into the right ventricular walls of rat hearts

for up to 28 days, showing improved neotissue formation in

terms of cell recruitment, proliferation, and blood vessel

density when compared to scaffolds without VEGF. A higher

density of VEGF immobilization was also shown to have a

greater resulting blood vessel density compared to lower VEGF

concentrations. In another study, Serpooshan et al. [72] used

compressed collagen type I as scaffolds for myocardial

infarction repair. While this is not a congenital disorder, this

patch may have applications for remodeling hearts with

tetralogy of Fallot. Their results after four weeks of

implantation exhibited limited fibrosis, diminished dilation of

the left ventricle, as well as neo-angiogenesis within the patch

when compared to the control.

Because decellularized matrices and purely collagen based

scaffolds both have disadvantages in creating a fully

functioning cardiac patch, other naturally derived polymers

such as alginate, chitosan, and silk fibroin continue to have

research interest [73]. Each of these polymers is extracted

from living organisms and has shown to have applications for

tissue engineering. Studies have shown that alginate, a natural

polymer found in cell walls of seaweed [74], combined with

RGD peptide can be formed into scaffolds for cardiac tissue

engineering. The RGD is important in cellular attachment to

the scaffold, and is immobilized by carbodiimide chemistry.

Shachar et al. [75] show the maintenance of key cardiac

markers, α-actinin, N-cadherin, and connexin-43, suggesting

that alginate-RGD immobilized scaffolds alleviate the need for

the addition of ECM proteins or Matrigel into scaffolds.

Chitosan is also a promising natural polymer investigated as a




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component for cardiac patch tissue engineering. It is a linear

polysaccharide derived from chitin, which is found in shellfish

exoskeletons. Alone, chitosan has been successfully applied to

the wound healing market as a clotting agent accelerating

hemostasis [76]. In tissue engineering, when chitosan is

combined with collagen, the resulting material has mechanical

properties superior than collagen alone. Kathuria et al. [77]

developed and characterized an elastic chitosan-gelatin cryogel

that could be used for potential tissue engineering applications.

This material was able to withstand cyclic deformations up to

40% without significant deformation with a Young’s modulus

ranging from 36-39 kPa.

Silk fibroin is investigated in a number of tissue engineering

applications because it is mechanically strong, noncytotoxic,

presents low immunogenicity, and is biodegradable. An

investigation into silk fibroin from A. mylitta silk worms shows

better cardiomyocyte attachment and functional beating after

seeding for up to 20 days [78]. These results are superior than

similar scaffolds using silk fibroin derived from mulberry B.

mori silk worms. This is proposed to be a result of A. mylitta

having RGD domains. Chi et al, [79] investigated a

chitosan-hyaluronan/silk fibroin cardiac patch implanted into

the left ventricles of rats. This patch was developed creating an

aqueous silk fibroin/chitosan/hyaluronan solution in a 10:1:1

ratio, and then spray-dried into patch form. Their implanted

patch showed reduced dilation of left ventricular diameter (4.27

0.29mm), increased wall thickness (1.5 0.13mm), and

improved left ventricular fractional shortening (42.8 2.4%).

Synthetic based cardiac patches

The use of synthetic polymers in surgical correction of CHDs

is limited to bioinert materials which can often illicit an

inflammatory response and fibrosis. The appeal of synthetic

polymers stem from their tunable mechanical, structural and

degradation properties. Fabrication of synthetic polymer based

scaffolds is very diverse and can involve techniques such as UV

polymerization [80], electrospinning [81], or laser sintering

[82], to name a few. Polymers such as polyethylene

terephthalate and expanded polytetrafluoroethlene (ePTFE)

have been used as cardiac patch materials in areas of lower

cyclic mechanical stress such as the septal wall. However both

polymers are not biodegradable and present issues in terms of

tissue ingrowth and remodeling. Ideally, corrections for CHDs

will only require one surgical procedure. In the case of cardiac

patches, this requires a delicate balance of new tissue ingrowth

aligned with scaffold degradation. Therefore much research

has focused on using biodegradable synthetic polymers such as

poly(lactic-co-glycolic acid) (PLGA), poly(plycerol sebacate)

(PGS), polyurethanes (PU), and polycaprolactone (PCL) as

scaffolds for cardiac tissue applications. The main issue when

using synthetic polymers as a scaffold material is inferior

cellular adhesion compared to that of natural polymers.

Therefore some strategies involve coating synthetic polymers

with other naturally derived materials, as discussed later in the

chapter.

Dacron® (polyethylene terephthalate) grafts are sometimes

used as septal defect patch material. Dacron® is a strong,

stable polymer that exhibits minimal degradation in vivo. Once

implanted, Dacron® elicits an inflammatory reaction and

subsequent fibrosis occurs. This is a structural solution for

defective cardiac tissue as applied to septal repairs, but the

material remains inert [83]. Another synthetic polymer,

ePTFE, is also currently used in repairing CHDs. The patch

materials developed from ePTFE are arranged spatially to have

pores ranging from 20-30 μm, which has shown to inhibit

cellular ingrowth [84]. Also, ePTFE does not induce as much

of a fibrous reaction as Dacron®, and therefore can be used as a

patch in areas of blood flow, such as a right ventricular outflow

reconstruction [85]. Both of these patch materials are far from

an ideal engineered cardiac tissue. Since they are not

remodeled and incorporated into existing tissue and elicit an

inflammatory response [86], they result in deficient mechanical

properties as well as hemodynamic changes. These polymers

will not degrade after implantation and will not promote

complete tissue remodeling. This is important in pediatric

patients that will continue to grow and therefore an inert patch

will likely require subsequent surgical operations.

Creating a synthetic material that is biocompatible and is

able to grow and remodel with a patient is no simple task.

Therefore bioresorbable polymer research into a viable patch

for CHDs has much attention. Like natural polymers, synthetic

bioresorbable polymers will be degraded over time in vivo

resulting from hydrolysis or enzymatic cleavage. The ideal

construct will provide mechanical support long enough for

native or seeded cells to produce and remodel ECM as well as

promote angiogenesis. In CHD patients, this allows for

continued growth without the need for additional surgical

interventions. Poly(lactic-co-glycolic acid) (PLGA) is a

biodegradable polymer that is used in many applications in

tissue engineering from drug delivery to hard and soft tissue.

The appeal of PLGA stems from the naturally metabolized

degradation products of lactic and glycolic acid. Zhou et al.

[55] showed that a PLGA scaffold wrapped with omentum

significantly improved ventricular remodeling and cardiac

function.

PGS has been studied frequently in the field of soft tissue

engineering. PGS is an appealing polymer that is able to

sustain and recover from deformation with minimal loss of

elasticity. Therefore the viscoelastic properties of PGS are

suggested to fit a mechanically dynamic environment such as

the heart. Chen et al. [87] developed a cardiac patch using

preconditioned PGS scaffolds in combination with human

embryonic stem cell-derived cardiomyocytes. The PGS

scaffolds were preconditioned for 6 days in media and

cardiomyocytes were seeded and shown to attach without the

need for a gelatin coating. The conditioned PGS scaffold was

able to sustain beating cardiomyocytes in vitro for longer than 3

months. Recently, Rai et al. [88] created a biomimetic PGS

scaffold attempting to chemically modify the surface of the

material in order to enhance cardiomyocyte attachment. This

was done by a process involving alkaline hydrolysis and

acidification to expose surface carboxyl chemical groups.

These groups could then be functionalized with peptides

YIGSR and GRGDSP in order to promote cellular attachment.

Their results showed that a ligand surface concentration of

10-15 mL/cm2 was sufficient to support attachment and growth

of both rat and human cardiac progenitor cells. The Karp lab

has expanded on the mechanical properties of PGS, creating




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poly(plycerol sebacate urethane) (PGSU). This co-polymer

blend was shown to have tunable mechanics and the ability to

deliver localized biomolecules when tested as a cardiac patch

material [89].

Fujimoto et al. [84] developed a polyester urethane urea

(PEUU) cardiac patch and implanted them into rats with

infarcted left ventricular wall. Their results suggested that the

scaffold promoted ingrowth of smooth muscle bundles with

mature contractile phenotype. Also, 8 weeks after implantation

the PEUU patch was largely resorbed, suggesting cellular

migration and improved cardiac remodeling.

PCL is another synthetic polymer of interest when

considering cardiac patch development. PCL is a

biodegradable polymer and can be hydrolyzed at the ester

linkages forming nontoxic byproducts. Yeong, et al. [82]

fabricated a PCL scaffold using a computer-aided selective

laser sintering technique. This fabrication technique showed

proof of concept for a customizable scaffold design allowing

for uniform control of pore size and patterning. By controlling

the laser wattage and scanning speed, PCL particles were

sintered into a disc shaped scaffold composed of small

repeating square pyramid units. In this particular design, they

were able to obtain a compressive stiffness of 345 kPa. Further

optimization of scaffold design could lead to tensile strengths

closer to native myocardium, which is on the range of 3-15 kPa.

It is important for a patch to exhibit stiffness similar to native

tissue as very elastic materials could form aneurysms and less

elastic materials could present with high local stress areas [90].

Natural/synthetic hybrid based patches

In order to take advantage of the cellular attachment and

signaling capabilities of natural polymers while maintaining a

way to control and optimize mechanical properties, researchers

have investigated a number of materials combining both natural

and synthetic materials. Pok et al. [90, 91] developed a

multilayered scaffold composed of PCL, chitosan and gelatin.

This scaffold was self-assembled with the stronger PCL core

sandwiched between emulsified solutions of gelatin/chitosan

(Figure 1). By controlling the average molecular weight of the

PCL core, the ultimate tensile strength of the scaffold could be

controlled in the range of 2-4 MPa. When combined with the

gelatin/chitosan, the compressive modulus of the scaffold was

close to native cardiac tissue (~15 kPa). This provided for a

biocompatible cardiac patch that when seeded with neonatal rat

ventricular myocytes resulted in spontaneous beating in a 50

vol.% gelatin:50 vol.% chitosan blend. Aside from layering

different polymers, the blending of both natural and synthetic

polymers without any chemical linkages show potential in the

development of viable cardiac scaffolds. Kharaziha et al. [92]

developed a scaffold composed of electrospun PGS and gelatin

nanofibers, showing that a 33 wt.% PGS formulation induced

optimal synchronous contractions of seeded cardiomyocytes.

Recent studies by Martins et al. [93] combine chitosan and

carbon nanofibers to create scaffolds that can enhance the

electrical properties of a patch material as well, showing a

tissue where conductivity is necessary to transmit signals for

beating. Being able to optimize the properties of both natural

and synthetic materials suggests a more adaptive and

structurally sound scaffold compared to a homogenous

construct.

Figure 1: Composite hydrogel patch with a multilayered

structure. The scaffold is made of a self-assembled PCL core

placed between gelatin–chitosan hydrogel layers. (A)

Macroscopic structure after disk formation. (B) Scanning

electron micrograph depicting a cross-section of the scaffold.

Reproduced from [90] with permission.

Injectable cardiac gels

Another highly researched area is the use of injectable

biomaterials for the treatment of myocardial defects. The idea

is that an amorphous matrix composed of critical tissue forming

components (cells, ECM proteins, small molecules) can be

injected directly to the site of deficiency, aiming to preserve

and promote cardiac tissue remodeling [94]. This strategy is

very appealing to myocardial infarction cases where the rapid

degradation of cardiac tissue is unrepairable. While CHDs do

not include myocardial infarction, a number of diseases

(coarctation of the aorta, anomalous origin of the left coronary

artery arising from the pulmonary artery, etc) [95] run the risk

of creating necrotic regions of cardiac tissue due to poor

perfusion or inadequate oxygen saturation of the systemic

blood [96]. As with myocardial infarction, the lack of

oxygenated blood due to CHD can manifest itself immediately

after birth and upon closure of the ductus arteriosus. Aside

from maintaining a PDA by administration of prostaglandin E1,

the use of an injectable cardiac matrix in CHD patients could at

a minimum help to preserve cardiac tissue until other corrective

action can be considered.

Similar to the cardiac patch studies, the variety of

biomaterials used for injectable gels can be both natural and

synthetic. Injectable gels are typically based on natural

materials such as fibrin, collagen, alginate, but there are some

gels which incorporate synthetic polymers either as hydrogel

copolymers or encapsulation vessels [97, 98]. Upon injection

into physiological temperatures (37oC), the gelling of the

components can occur very rapidly, within seconds in most

cases. The other key component of injectable materials are

living cells aimed to assist in regenerating damaged tissue by

paracrine delivery of different signaling molecules such as

growth factors, cytokines, hormones, etc. The key advantages

to an injectable strategy are: 1) they are minimally invasive and

easy to administer through injection, 2) amorphous structure

allows for high contourability allowing the gel to fill various

defect shapes and sizes and 3) therapeutic agents are easily

incorporated and delivered to the defect site.

Fibrin glue has been studied as an injectable material for

myocardium repair. Fibrin glue is currently FDA approved as a




8

hemostasis sealant for use during surgical operations [99].

When applied as an endoventricular heart patch, Christman et

al. [100] showed that the use of a fibrin glue increased cell

transport and survival of skeletal myoblasts, decreased the size

of infarcted left ventricle, and increased blow flow to the area

of myocardial ischemia compared to controls. Other studies

focused on the delivery of growth factors with the goal of

enhancing native cardiac cells to survive and proliferate. The

Christman group also developed an injectable material

composed of decellularized heart matrix [101], which was

shown to have endothelial and smooth muscle cell migration as

well as arteriole formation after 11 days. Ruvinov et al. [102]

show that an alginate based gel can be used to sequentially

deliver insulin-like growth factor (IGF-1) and hepatocyte

growth factor (HGF) aimed to induce myocardial regeneration.

The alginate gel sufficiently prevented proteolysis of the two

proteins and when injected into a rat acute myocardial

infarction model, infarct expansion was attenuated and

increased angiogenesis throughout affected area was observed

after 4 weeks. A combination of growth factors and stem cells

as a regenerative therapy is popular as well. This concept will

not only deliver signaling molecules to native cardiac cells, but

provide additional cells that may differentiate and aid in the

remodeling process. Wang et al. [97] developed an injectable

hydrogel composed of collagen type I, chondroitin sulfate, and

a thermosensitive copolymer (Figure 2). This hydrogel was

shown in vitro to be capable of releasing IGF-1 over a 2 week

period in order to enhance the survival and growth of

encapsulated mesenchymal stromal cells (MSC). The

differentiation potential of the MSCs was maintained within the

hydrogel, but the addition of IGF-1 was shown to significantly

accelerate MSC growth. The Davies group conducted studies

into the effects of injectable gels on cardiac tissue remodeling,

showing that injection of PEG based gels following infarction

could aid in the immediate healing and remodeling [103]. They

also showed a temporal remodeling relationship, where the

same gels were injected one week post-infarction had slower

degradation rates compared to gels injected immediately

post-infarction [104].

The major disadvantage of injectable biomaterials for the

repair of infarcted cardiac tissue is that they lack sufficient

stiffness compared to the tissue in chronic diseased states.

Where healthy adult myocardium has a modulus of

approximately 50 kPa , diseased states can range from 200-300

kPa [105]. Ifkovits et. al. elucidated the effects of stiffness on

ovine infarct models using injectable methacrylated hyaluronic

acid hydrogels with varying moduli. The study shows that a

higher modulus (~43 kPa) gel applied post-infarction resulted

in less infarct expansion and reduced left ventricular dilation

compared to a lower modulus (8 kPa) [106]. While these

injectable gels nearly match the modulus of healthy adult

myocardium, the stiffness may not be sufficient to support

chronic diseased myocardium. However, injectable

biomaterials may have potential in pediatric cases where years

of cardiac remodeling have not altered the mechanical

properties of the myocardium.

Figure 2: An injectable, flexible, thermally responsive

hydrogel composite for cardiovascular tissue engineering

applications. The material is comprised of type I collagen,

chondroitin sulfate, and a thermosensitive copolymer based on

N-isopropylacrylamide. (a) The composite is liquid at 4°C. (b)

Hydrogel forms upon temperature increase to 37°C. (c) The

material is injectable through a 26-gauge needle at 4°C. (d)

Hydrogel strip at 37°C before stretching. (e) Hydrogel strip at

37°C after stretching. Reproduced from [97] with permission.

4. HEART VALVE TISSUE ENGINEERING IN CONGENITAL HEART

DISEASE

The direction of blood flow through the heart is regulated by

the heart valves (aortic, pulmonary, mitral and tricuspid valve).

All four heart valves can be affected by a congenital defect.

Problems include stenosis and regurgitation, and both cases

impede physiological blood flow. Ebstein’s anomaly is

characterized by an abnormal structure of the tricuspid valve

and regurgitation. Another type of congenital heart defect is a

bicuspid aortic valve where two cusps are present instead of

three. This defect does not usually involve severe

complications in the pediatric population and surgical

correction, if deemed necessary, is performed well into

adulthood [107]. Anomalies of the pulmonary valve are the

most common among the four valves and are often manifested

in conjunction with other congenital heart defects [108]. Also,

in certain lesions, the pulmonary valve may be underdeveloped

and a reconstruction of the right ventricular outflow tract

(RVOT) is a beneficial option. A valved conduit is used to

connect the right ventricle to the pulmonary artery [109].

When replacing a failing heart valve, a surgeon generally

has the following options: A mechanical valve, or a

bioprosthetic valve which can be either a homograft or a

xenograft. The use of a mechanical valve is associated with

increased risk for thrombosis and patients will need lifelong

anticoagulant medication. Anticoagulants can prove dangerous

for children due to possible hemorrhaging which is a threat in

this physically active patient population. Therefore,

bioprosthetic valves are usually preferred for surgical

correction in pediatric patients. Human donor heart valves are

commonly cryopreserved until implantation, [110] but a main

drawback is immunogenicity [111]. Decellularization

techniques have been more recently proposed to reduce this

effect in homografts [112]. As with all homografts, availability




9

is a common limiting factor. In the pediatric population, this

proves even more problematic as a valve of the right size needs

to be found and transplanted for each child. Xenograft valves

are typically bovine or porcine tissues crosslinked with

glutaraldehyde. While xenografts are more readily available

than homografts, the concern about immunogenicity remains

[113, 114]. Bioprosthetic valves, both human and

animal-derived, oftentimes suffer from early degeneration

which can be detrimental to young patients, necessitating

replacement operations. One of the mechanisms of

bioprosthetic heart valve degeneration that can lead to failure is

linked to calcification. The phenomenon of bioprosthetic heart

valve calcification has been reviewed elsewhere [115].

As an alternative to tissue crosslinking, decellularization

has been introduced in order to preserve the original tissue

matrix to a certain extent while reducing cellular antigens.

Commercially available decellularized heart valve porcine

xenografts (e.g. Synergraft®, Matrix P®) and homografts

(CryoValve SG®) exist that are suitable for use in a pediatric

congenital heart disease population. However, failure resulting

in death or need for graft replacement has been reported for

porcine decellularized heart valves due to a possible immune

response and limited cell migration from the host causing poor

recellularization [116-118].

Another option for the surgical management of aortic valve

disease in children is the Ross procedure. During this

procedure, the aortic valve is replaced by the patient’s own

pulmonary valve. Since an autograft is placed in the aortic

position, the commonly associated problems with mechanical,

homo-or xenograft valves are eliminated, but it generates the

need for a suitable replacement for the pulmonary valve [119].

The autologously implanted aortic valve has been shown to

grow and function satisfactorily in young children [120].

For all types of currently available options, excluding the

new aortic valve after the Ross procedure, as the child grows,

reoperation will be necessary to provide a new valve that

matches their size. Additionally, the prosthetic tissues have

limited intrinsic potential for repair. General risk factors for

prosthetic heart valve complications in children have been

identified, which include age and type of defect [110, 121].

Therefore, the field of tissue engineering has been faced with

much anticipation for a solution that can eliminate some of the

current limitations in heart valve replacement for children.

Heart valves are unique tissues considering their

architecture and the biomechanical environment in which they

operate. Even though all four valves show a similar pattern in

cell phenotypes and extracellular matrix components, they have

structural and morphological variances. These differences in

structure allow valves to optimally serve their function, for

example the high pressure setting of the aortic valve presents

different mechanical loading compared to the lower-pressure

pulmonary circulation. Moreover, heterogeneity and anisotropy

is pronounced within regions of each valve [122-124].

Therefore, a profound understanding of the properties of each

valve is important in order to create tissue engineered analogs.

The field of heart valve tissue engineering is largely benefitting

from research performed in the biomechanics, mechanobiology

and cell biology of heart valves [125-128].

Approaches involving naturally derived matrices

Building upon the existing surgical practice of bioprosthetic

valves, heart valves have been decellularized and reseeded with

cells following a tissue engineering paradigm. The goal of this

effort is to combine the advantageous properties of a native

matrix with the regenerative and remodeling potential of living,

autologous cells. The first clinical case of this kind was

described in 2002. A cryopreserved pulmonary homograft was

decellularized and reseeded with endothelial cells from the

patient's forearm vein. One year after implantation, the patient

was in good clinical condition with appropriate valve function

[129]. The same group reported the results of clinical studies

performed over ten years. Eleven adult patients, some of them

with congenital heart disease, underwent the Ross procedure

and a tissue engineered heart valve (TEHV) was placed in the

pulmonary position. The TEHV was fabricated as above. At the

ten-year follow-up, the survival rate was 100% and there were

no signs of valve regurgitation, calcification or morphologic

degradation. These results were encouraging, however more

long-term studies are needed to evaluate the longevity of these

valves, especially in a younger patient population [130]. A

clinical report involving the placement of decellularized and

reseeded pulmonary valves in two pediatric tetralogy of Fallot

patients was reported in 2006. Autologous endothelial

progenitor cells were harvested from peripheral blood and

seeded on valve homografts. Both children were doing well

clinically at the 3.5 year follow-up. The valves showed

sufficient mobility, no degeneration or stenosis, and

importantly, they were shown to grow with the children [131].

These approaches have shown promising clinical value, but

certain limitations are still present. The decellularized matrix is

a homograft, and the donor scarcity problem remains. The cell

harvesting, often done in an invasive manner, as well as the

isolation and expansion procedures, impose an additional step

which is time-and effort-consuming. Adequate quality control

systems must be in place. Finally, there are questions about the

fate of the seeded cells and their role in the regenerative

process. Work by Roh et al. [132] has shed some light on the

last question and will be discussed in the next chapter.

Therefore, some researchers have focused on the in situ

recellularization from the host. This approach falls under the

classical tissue engineering paradigm, which involves

scaffolds, cells and stimuli, and where cell seeding happens in

vivo. Commercially available, decellularized and not re-seeded

matrices have not always been very successful clinically, and

even devastating in some cases, as previously mentioned

[116-118]. Factors that will stimulate and improve the active in

vivo recellularization are currently being investigated. Jordan et

al. decellularized porcine pulmonary valves and conjugated

them with an antibody that recognized ovine endothelial

progenitor cell (EPC) markers. The valves were implanted in

sheep in the pulmonary position. At three months

post-implantation, histological characterization showed

significantly higher cell numbers in the antibody-conjugated

valves compared to unconjugated, unseeded matrices or valves

seeded with EPCs that served as controls. On the latter type, the

initially seeded cells seemed to have disappeared within the

first month (Figure 3). The mechanical properties of the

conjugated valves were also found to be superior compared to




10

the controls [133]. Moreover, in vitro studies have shown that

the anisotropic characteristics of the valve microenvironment

have an effect on cell migration and subsequent

recellularization [134]. These findings may prove useful

towards the rational design of tissue engineering strategies for

the in situ cellular repopulation of a previously decellularized

matrix.

Figure 3: Porcine pulmonary valves were decellularized and

either (1) not treated (un-conjugated), (2) reseeded with

autologous endothelial progenitor cells (cell-seeded), or (3)

conjugated with a CD133 antibody specific for ovine

endothelial progenitor cell markers (conjugated). A)

Hematoxylin and eosin stained sections of valve leaflets after 3

months in vivo. Unconjugated and cell-seeded valves show

very few cells. Conjugated valves showed staining indicating

cell content throughout the valve. B) Cell content in conjugated

valve leaflets after 1 week, 1 month, and 3 months in vivo.

Increased recellularization is observed over time with

4,6-diamino-2-phenylindole nuclear staining. Reproduced from

[133] with permission.

Scaffolds based on natural polymers have also been used in

heart valve tissue engineering, albeit not yet on a clinical level.

Collagen and glycosaminoglycans (GAG) are two of the main

ECM components in all four native heart valves [135], and

efforts have been directed towards their incorporation into

tissue engineering. Collagen can be obtained from animal

tissues which are in adequate supply. Tedder et al. have

prepared collagen scaffolds from decellularized porcine

pericardium and pulmonary arteries. Layered constructs

incorporating collagen derived from those two tissues could be

shaped into a heart valve with unique properties in each region.

They have also suggested the use of penta-galloyl glucose, a

collagen-binding agent, as an alternative to glutaraldehyde in

order to stabilize collagen from premature in vivo degradation

without inducing calcification [136, 137]. Collagen hydrogels

have also been rendered more porous and bioactive with the

addition of GAG, thereby supporting co-cultures of valvular

interstitial cells (VIC) and endothelial cells [138]. Fibrin is

another protein introduced to heart valve tissue engineering. It

can lead to a completely autologous approach as this protein

can be isolated from the patient’s own blood. Studies have

shown that fibrin can be used in a mechanical loading

environment and promote ECM production when seeded with

cells [139-141]. Hyaluronic acid (HA) is a polysaccharide and a

major GAG component of the heart valve extracellular matrix.

Early work demonstrated the potential of crosslinked HA

hydrogels as cell culture vehicles and assessed proliferation and

ECM production [142, 143]. The molecular weight and degree

of modification of HA has an effect of the resulting hydrogel

properties, which in turn influence VIC phenotype and

behavior [144]. Interactions between VIC and HA have been

shown to have implications in mineralization [145]. This

observation can be useful for HA-based tissue engineering

strategies towards the prevention of heart valve calcification.

Approaches involving synthetic matrices

Biodegradable polyesters such as poly(glycolic acid) (PGA)

and poly(lactic-co-glycolic acid) (PLGA) have been some of

the first synthetic materials introduced to heart valve tissue

engineering [146]. PGA, PLLA and PLGA scaffolds are still

widely used by many investigators. The group of M. Sacks has

been examining the effect of biomechanical stimuli on cells

seeded on these types of materials towards the development of

a heart valve tissue engineering equivalent. They recently

reported the effects of physiological conditioning of tri-leaflet

heart valve constructs in a bioreactor in which oscillatory

pressure and flow conditions could be adjusted to simulate

pulmonary artery hemodynamic conditions. This bioreactor

allows for investigation of the effects of organ level, dynamic

pulsatile culture conditions such as mean pressure, mean flow

rate, beat frequency, stroke volume and shear stress [147].

Increased collagen production was observed in the

biomechanical environment of an organ level, tri-leaflet valve

as compared with a bioreactor system that applies flexure and

flow-induced shear stress to rectangular, cell-seeded scaffolds

[148]. They and others [149] have stressed the importance of

biomechanical stimuli for optimal tissue engineered heart valve

formation.

Poly(glycolic acid) mesh scaffolds coated with

poly(4-hydroxybutyrate) (PGA/P4HB) have been used

extensively by Hoerstrup and collaborators over the past

decade. P4HB is a naturally derived biopolymer which is

biodegradable and pliable. They first reported seeding of these

scaffolds with myofibroblasts and endothelial cells and

culturing them in a dynamic bioreactor system before

implantation in the pulmonary position in sheep. After twenty

weeks in vivo, the scaffold material was completely degraded

and the neo-tissue was similar in architecture and mechanical

properties to native heart valve leaflets [150]. More recently,

the group reported use of these scaffolds with amniotic fluid

derived stem cells (AFSC). These cells could be harvested

prenatally following the detection of a heart defect in a fetus,

and a tissue engineered construct could be fabricated in vitro

and be ready for implantation by the time of birth. AFSC




11

showed growth and differentiation potential when cultured in a

bioreactor system. The mechanical properties of the resulting

valves did not reach physiological values, but sufficient

opening and closing of the valves was observed [151]. Seeded

scaffolds were implanted in sheep fetuses in the pulmonary

position via a minimally invasive transapical technique (Figure

4). This work served as a first proof of concept for a single-step,

fetal cardiac interventional procedure with potential

implications in congenital heart disease [152]. Another

approach presented by the same group involved the

decellularization of a PGA/P4HB scaffold that was previously

seeded with human myofibroblasts and cultured dynamically in

a bioreactor. These decellularized tissue engineered heart

valves showed significantly higher recellularization in vivo in a

non-human primate model as compared to decellularized native

homograft valves, which served as controls. The authors

suggest that this in situ tissue engineering approach might

facilitate the creation of “off the shelf” non-immunogenic heart

valve analogs [153].

Figure 4: Amniotic fluid stem cell-based heart valves. (a)

Matrices were fabricated from PGA and placed within

self-expandable stents. Asterisk indicates the three valve

leaflets. (b) Matrices were coated with P4HB. Demonstrated

here is the micro-CT image with arrows indicating the

PGA-P4HB-composite matrix. (c-d) Amniotic fluid stem cells

were seeded onto the valves using fibrin as a carrier. Grating

interferometry showed homogeneous cell seeding throughout

the scaffold. Arrows indicate fibrin matrix. (e) Macroscopic

view of cell-seeded matrix. (f-h) The cell-seeded matrices were

crimped and placed inside the delivery system. (i) Planar

fluorescence reflectance imaging was performed to analyze cell

loss during crimping. Red signal indicates autofluorescence

and yellow is specific signal for CFSE-labeled cells. (j)

Confocal microscopy image of cell-seeded matrix.

CFSE-labeled cells appear green. Reproduced from [152] with

permission.

Other synthetic materials currently used in heart valve tissue

engineering include polycaprolactone [154], poly(glycerol

sebacate) [155] or combinations thereof [156]. Polyurethanes

have been also attractive because of their largely tunable

properties [157, 158]. Poly (ethylene glycol) (PEG) hydrogels

have been used as three-dimensional matrices for valvular

interstitial cell encapsulation. Studies have led to a better

understanding of the cells’ mechanobiology while suggesting

that PEG-based hydrogels can be tailored to provide an

instructive environment for the development of more mature

tissues [159, 160]. Hydrogels also allow for the fabrication of

layered constructs which can mimic the architecture and

mechanical function of each region of the native valve.

Tri-layered PEG hydrogels with two stiff outer layers and a soft

inner layer as proposed by the researchers have potential of

creating a multilaminate environment for cell encapsulation

[161]. In another biomimetic approach, PEG-PLLA hybrid

scaffolds were electrospun and each region was functionalized

with biomolecules present in the extracellular matrix of native

heart valves [162]. In summary, promising results have been

obtained with synthetic matrices, however a better

understanding of their in vivo behavior, especially with regard

to cell infiltration and biomechanical performance, is necessary

before clinical translation.

5. BLOOD VESSEL TISSUE ENGINEERING IN CONGENITAL HEART

DISEASE

It is important to differentiate how disorders involving the

blood vessels which may benefit from tissue engineering

present in children as compared to adults. Whereas heart

disease affecting children is mostly congenital, many forms of

acquired heart disease in adults stem from atherosclerosis,

which is caused by the buildup of fatty plaque inside the

arteries. This results in narrowing of the arteries leading to a

reduction in tissue perfusion. Coronary artery disease is a

common and serious manifestation. Bypass surgery is one of

the possible treatments for coronary artery disease, in which a

vascular graft is required.

Congenital heart defects involving blood vessels include

coronary artery anomalies [163], coarctation (narrowing) of the

aorta [164], interrupted aortic arch [165], and pulmonary artery

stenosis [166]. However, there are certain cases of complex

congenital heart disease, in which the blood flow to and from

the heart needs to be rearranged. This rearrangement allows for

proper oxygenation of the blood and improvement of

hemodynamic properties. In a structurally normal, biventricular

heart, the systemic and pulmonary circulations are each

supported by a ventricle and the circulations are in series. One

particular congenital defect is a heart with only one functional

ventricle. In patients born with a single ventricular chamber, the

two circulations are in parallel and patients only survive

because the systemic and pulmonary venous bloods mix [167].

A palliative surgical procedure that has been applied for

treatment of children born with a functional single ventricle is

the Fontan operation and its variations. The various

Fontan-type procedures involve diverting the systemic venous

blood to the pulmonary arteries, bypassing the right heart [168,




12

169]. These surgeries are performed in stages and grafts are

required for the vascular reconstruction.

Solutions used by pediatric cardiothoracic surgeons include

synthetic and biological grafts. Synthetic polymer grafts such

as polyethylene terephthalate (Dacron®) and the currently

more prevalent expanded polytetrafluoroethylene (ePTFE,

Gore-Tex®) have been associated with increased risk for

thrombus formation, stenosis, foreign body reaction, and

calcification [170]. Due to the increased risk of thrombosis,

life-long use of anticoagulants is essential. Biological grafts

include auto-and homografts, which have a lower tendency for

thromboembolic events, but may show early degeneration due

to increased calcification [171, 172]. Additionally, their supply

is not always guaranteed. Vascular grafts presently employed

for Fontan-type operations and their limitations are

summarized in a review article by Patterson et al [173]. For

pediatric surgery, a vascular graft typically has smaller

diameters than those needed for similar procedures in adults,

and most importantly, it needs the ability to grow with the

patient and have high life expectancy. One exception where

growth is not required is the case of vascular grafts used in

temporary shunt systems as for example the modified

Blalock-Taussig shunt. This shunt connects the systemic to the

pulmonary artery and provides short-term palliation in complex

cyanotic cases until further corrective or palliative surgery can

be performed [174].

General considerations for blood vessel tissue engineering

Progress in the field of blood vessel tissue engineering has

brought some initial successes, and important basic research

has improved our understanding of the physiological

remodeling process. Some recent review articles [175-177]

summarize the advances in the general area of blood vessel

tissue engineering. As with other tissue engineering paradigms,

approaches include the use of cells with and without scaffolds,

which can be either decellularized native matrix, or a

natural/synthetic material. Some of the remaining challenges to

be resolved before tissue engineered blood vessels can be

translated into clinical practice have been in highlighted in

articles by L’Heureux et al. [178] and Udelsman et al. [179].

In the case of cell-seeded matrices, the cell source must be

easily accessible and cells should be readily expandable in

sufficient numbers in order not to slow down the

transplantation process. Additionally, a reproducible cell

seeding technique is important [180, 181]. Endothelial cells

play an important role in the function of blood vessels. They

serve as the interface between the blood and the vessel wall,

and among other functions, increase thromboresistance. The

presence of endothelial cells in tissue engineered blood vessels

has been shown to impart patency and prevent the formation of

thrombi [182]. Early studies by Campbell et al. [183] utilized a

silastic tube placed in the peritoneal cavity of rats and rabbits

which was covered in granulation tissue after two weeks in

vivo. This tissue was removed from the silastic tube, which

served as the inflammatory agent, and inverted. A layered

structure with myofibroblasts and collagen was observed.

Mesothelial cells, or migrated endothelial cells, which stained

positive for von Willebrand factor, formed the innermost layer.

The tube made of tissue was implanted in experimental animals

and served as a patent vascular graft over a period of four

months. In recent years, the quest for the elucidation of the

mechanism of endothelialization in vivo along with the

questions on graft stenosis as observed in early clinical trials

[184] have led to the intense investigation of the molecular and

cellular processes that govern the in vivo remodeling of tissue

engineered blood vessels. It was demonstrated that scaffolds

seeded with bone marrow mononuclear cells elicited an

inflammatory response in vivo, which attracted host

monocytes. The graft matured into a living blood vessel

repopulated with endothelial and smooth muscle cells over six

months. Interestingly, the initially seeded bone marrow

mononuclear cells were not detectable a few days after

implantation [132]. Further studies confirmed the adjacent

blood vessel wall as the source of the endothelial and smooth

muscle cells that were found in the tissue engineered blood

vessel graft [185]. Moreover, macrophage infiltration was

shown to be involved in the formation of stenosis in a murine

model [186]. These studies proved the importance of the

inflammation process in the in vivo remodeling of tissue

engineered blood vessels, and have initiated a discussion on the

role of seeded cells in the various tissue engineering models.

Blood vessel tissue engineering in congenital heart disease:

First clinically relevant attempts

Besides the need for small diameter (<6mm) vascular grafts

for coronary artery bypass surgeries in adults, for which no

ideal and readily available substitute exists, tissue engineering

has been also recognized as a potential solution for pediatric

vascular grafts. One of the first reported studies geared

specifically towards congenital heart defects was the work on

tissue engineered pulmonary artery autografts in 1998. Shinoka

et al [187] performed an in vivo study in sheep replacing a 2-cm

segment of their pulmonary arteries with tissue engineered

grafts. Tubular scaffolds were made of biodegradable

polyglactin/ polyglycolic acid mesh. The cellular source was

autologous artery or vein. Cells were separated and the

endothelial cell-rich fraction was seeded in the lumen, whereas

the endothelial-deficient fraction was seeded on the periphery

of the scaffold. After implantation, the sheep were monitored

by echocardiography and angiocardiography, and samples were

harvested in intervals up to six months. No scaffold material

was detectable after 11 weeks. Seeded scaffolds were patent

and did not show macroscopic signs of calcification, but

evidence of ECM remodeling and diameter increase was

observed.

The first successful clinical application of a tissue

engineered vascular graft was reported in 2001 in Japan. A

four-year old girl with single ventricle physiology received a

construct made of a PLLA/PCL copolymer reinforced with

PGA fibers and seeded with autologous peripheral vein cells.

The cells were previously expanded over a period of eight

weeks. The patient had previously undergone the Fontan

procedure and a 2-cm segment of her pulmonary artery was

occluded and needed replacement. The tissue engineered graft

had been cultured in vitro for ten days prior to implantation, and

measured 10 mm in diameter, 20 mm in length and 1 mm in

thickness. The polymer was designed to degrade within eight

weeks in vivo. No complications were reported immediately




13

after the surgery, and seven months post-operatively, the

patient was doing well [188]. Soon after, the same group

initiated a clinical trial in Japan involving a biodegradable

scaffold composed of PGA and PLLA or PCL formed into a

tubular graft. Autologous bone marrow-derived mononuclear

cells were obtained and seeded onto the scaffolds. The recipient

cohort comprised of twenty-five patients with single ventricle

physiology with an age range from 1 to 24 years. After the

surgery, patients were treated with anticoagulants for three to

six months and were monitored by angiography,

echocardiography and computed tomography. No adverse

reactions were reported immediately postoperatively. At a

late-term follow up (mean of 5.8 years after implantation) no

evidence of graft rupture, aneurysm, or ectopic calcification

was observed (Figure 5). No graft related mortality was

reported, however one patient died 6 months post implantation

of congestive heart failure, and three more within four years.

The later-term deaths were related to other cardiovascular

anomalies. Graft stenosis was observed in 24% of the patients.

This study as the first of its kind demonstrated the feasibility

and relevance of tissue engineered vascular grafts in congenital

heart surgery. It also raised questions about the in vivo

mechanisms of vessel remodeling and the causes of graft

stenosis [184].

Figure 5: Results from the first human trial involving tissue

engineered blood vessels in congenital heart disease. A tissue

engineered vascular graft was fabricated by seeding

mononuclear cells on a biodegradable scaffold and evaluated in

pediatric patients with single ventricle physiology. Image

depicts a computed tomography (CT) scan one year after the

tissue engineered vascular graft was implanted. The arrows

indicate the extracardiac graft. No signs of aneurysm formation

and graft rupture were observed, and the graft remained patent.

Reproduced from [184] with permission.

Further advances in blood vessel tissue engineering for

congenital heart disease

Hoerstrup et al. [189] demonstrated the fabrication of

pulmonary artery conduits using human umbilical cord cells

seeded on synthetic biodegradable PGA/P4HB scaffolds and

cultured in a bioreactor. Morphologic characterization of the

tissue engineered pulmonary artery showed ECM formation

and myofibroblast viability, and the construct’s mechanical

properties were comparable to human tissue. The authors

proposed umbilical cord cells as a new and easily available cell

source for tissue engineering with potential applications in

congenital heart disease. These cells could be harvested at birth

and cultured in vitro with promising growth potential. More

recently, the same authors evaluated amniotic-fluid derived

stem cells which is another attractive cell source for congenital

applications. They were able to construct in vitro small-and

large-diameter vascular grafts as a first pre-clinical attempt

[153].

Another approach suggested the use of a decellularized

allogeneic matrix as a scaffold for pulmonary artery tissue

engineering [190]. Ovine pulmonary arteries were

decellularized and seeded with autologous endothelial cells

obtained from carotid arteries, and subsequently implanted in

sheep. Unseeded, decellularized arteries served as controls.

After six months, no calcification or thrombus formation was

observed in any of the explanted blood vessels. All vessels

showed an increase in diameter; however the unseeded blood

vessels showed a disproportionate increase in diameter

resulting in aneurysm formation. This study served as a proof

of concept for the in vitro re-endothelialization of

decellularized pulmonary arteries as an alternative to

biodegradable synthetic scaffolds, and their short-term

performance in vivo.

The field of congenital heart disease may benefit from the

overall developments in the field of blood vessel tissue

engineering. Some challenges that remain are the translation

into high-pressure areas such as for congenital defects of the

aorta, as early studies by Shinoka and colleagues have been in

low-pressure systems. Compliance properties of the neo-vessel

should match those of the native vessels as the child grows.

Finally, more information on long-term outcomes is warranted.

The first clinical trial in the United States specifically geared

towards pediatric heart defects was initiated in 2009 with Dr.

Christopher Breuer as the principal investigator (A Pilot Study

Investigating the Clinical Use of Tissue Engineered Vascular

Grafts in Congenital Heart Surgery) [191]. Results of the study

had not been published at the time of writing this manuscript.

6. CONCLUSIONS

Current surgical management options for the treatment of

congenital heart disease have some limitations, and outcomes

could be potentially improved by employment of tissue

engineering strategies. Significant progress has been made in

the tissue engineering of cardiac patches, heart valves and

blood vessels, with valuable implications for their applicability

in congenital heart disease. Clinical studies with tissue

engineered heart valves and blood vessels have generated

excitement and paved the way for further research. However, a

clear appreciation of the in vivo processes is warranted for the




14

successful clinical translation of tissue engineering strategies.

Current research is broadening our knowledge about the in vivo

molecular mechanisms that govern cell behavior as well as

material-host interaction. Pediatric cardiac tissue engineering

shares most of the challenges that the general field is facing,

and some unique ones. A thorough understanding of the

patients and their needs is important so that appropriate

strategies can be designed. Ideally, a graft that grows with the

patients and one that is durable over their entire life span is

desired. In order to achieve this goal, more long term studies

need to be implemented. Appropriate animal models that can

parallel children’s growth and can demonstrate the long term

behavior of tissue engineered implants are necessary. Many of

the current studies are performed on adult animals, which may

not prove sufficient indicators for pediatric cardiac tissue

engineering. Scaffold materials might provoke varying

inflammatory responses and exhibit different degradation

profiles in younger versus adult bodies. The benefits of

including autologous cells in the tissue engineering paradigm

for children must be evaluated. Often, time is an extremely

critical factor and cell expansion can be time-consuming,

unless potent cells can be harvested prenatally in a safe manner

and be ready for use by the child’s birth. Lastly, the potential of

tissue engineering stems from its multidisciplinary nature.

Great progress can be achieved if research projects can bring

together basic scientists, engineers, and pediatric clinician

specialists, to work in a concerted effort. Once these challenges

are overcome, there is great opportunity for tissue engineering

to provide a solution to the large number of children affected by

congenital heart disease.

ACKNOWLEDGMENTS

The authors would like to thank Dr. Wilson Lam for

consultation on the congenital heart defects. Grant support

through NSF (CAREER to J.G.J.) and AHA (Beginning Grant

in Aid to J.G.J.) is gratefully acknowledged.

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