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Review ArticleVascular Tissue Engineering: Recent Advances inSmall Diameter Blood Vessel Regeneration

Valentina Catto,1,2 Silvia Farè,1,2 Giuliano Freddi,3 and Maria Cristina Tanzi1,2

1 Laboratorio Biomateriali, Dipartimento di Chimica, Materiali e Ingegneria Chimica “G. Natta”, Politecnico di Milano,Piazza L. da Vinci 32, 20133 Milano, Italy

2 Local Unit Politecnico di Milano, INSTM, Via G. Giusti 9, 50121 Firenze, Italy3 INNOVHUB-SSI, Divisione Stazione Sperimentale per la Seta, Via G. Colombo 83, 20133 Milano, Italy

Correspondence should be addressed to Silvia Fare; [email protected]

Received 29 October 2013; Accepted 21 November 2013; Published 27 January 2014

Academic Editors: B. Hambly and M. Zhang

Copyright © 2014 Valentina Catto et al.This is an open access article distributed under the Creative Commons Attribution License,which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

Cardiovascular diseases are the leading cause of mortality around the globe. The development of a functional and appropriatesubstitute for small diameter blood vessel replacement is still a challenge to overcome the main drawbacks of autograftsand the inadequate performances of synthetic prostheses made of polyethylene terephthalate (PET, Dacron) and expandedpolytetrafluoroethylene (ePTFE, Goretex). Therefore, vascular tissue engineering has become a promising approach for smalldiameter blood vessel regeneration as demonstrated by the increasing interest dedicated to this field. This review is focused on themost relevant and recent studies concerning vascular tissue engineering for small diameter blood vessel applications. Specifically,the present work reviews research on the development of tissue-engineered vascular grafts made of decellularized matrices andnatural and/or biodegradable synthetic polymers and their realization without scaffold.

1. Introduction

Cardiovascular diseases are the leading cause of death aroundthe world. In 2008, 17.3 million people died from cardio-vascular related reasons; specifically 7.3 million were due tocoronary heart disease [1]. Currently, in the United States17% of overall national health expenditures is linked withcardiovascular diseases [2].

Bypass surgeries are commonly performed to allow theperipheral or coronary revascularization. To date, autograftsremain the standard clinical approach for the replacementof small diameter blood vessels (inner diameter (ID) <6mm). Nevertheless, autografts (such as saphenous vein,arm vein, mammalian artery, or radial artery [3, 4]) showsome drawbacks: considerable morbidity associated withautologous harvest and scarce availability due to diseasesor previous organ harvesting [4, 5]. Arterial autografts aremore indicated for coronary by-pass surgeries due to theirhigher mechanical properties [6]; internal mammary artery

showed a higher patency rate than saphenous vein (85%versus 61%, after 10 years) [7]. Synthetic materials, forexample, polyethylene terephthalate (PET) and expandedpolytetrafluoroethylene (ePTFE), are successfully used for thereplacement of medium-large diameter blood vessels (ID >6mm), when high blood flow and low resistance conditionsprevail [3, 4, 8]. However, synthetic grafts used for below-the-knee vascular by-pass and coronary by-pass (ID < 6mm)fail for unacceptable patency rates in the long term. Patencyof ePTFE prostheses is 40–50% when used to bypass theproximal popliteal artery at 5 years and 20% when usedfor infrapopliteal bypass at 3 years [3]. The use of PETor ePTFE for small diameter blood vessels leads to severalcomplications like aneurysm, intimal hyperplasia, calcifica-tion, thrombosis, infection, and lack of growth potential forpediatric applications [3, 4, 9]. These drawbacks are mainlycorrelated to the regeneration of a nonfunctional endothe-lium and a mismatch between the mechanical propertiesof grafts and native blood vessels [3, 5, 6]. A comparison

Hindawi Publishing CorporationISRN Vascular MedicineVolume 2014, Article ID 923030, 27 pageshttp://dx.doi.org/10.1155/2014/923030

2 ISRN Vascular Medicine

between the compliance and the elastic modulus of syntheticprostheses and those of native vessels is reported in Table 1.

Causes of graft failure may be classified into early,midterm, and late [10, 11]. Early failures (within 30 days afterthe implantation) are related to technical complications, flowdisturbances, or acute thrombosis [10, 11]. Midterm failures(3 months to 2 years after the implantation) consist of lumenocclusion due to intimal hyperplasia, while late failures (>2years) are related to atherosclerotic disease [10, 11].

With the aim to improve the scarce patency of syntheticgrafts, Deutsch et al. [12] developed a procedure for the autol-ogous in vitro endothelialization of ePTFE prostheses (ID =6-7mm). Specifically, the inner side of ePTFE prostheses wascoated with fibrin to allow for seeding of autologous ECs in invitro rotating conditions; after about 9 days of in vitro culture,341 grafts were implanted as infrainguinal bypass in 310patients [12]. During 15 years of clinical use, endothelializedePTFE prostheses showed the presence of endothelium 2–4years after the surgery and a patency rate similar to vein grafts[12]. The patency rate of 7mm prostheses was significantlyhigher than that of 6mm grafts (78% versus 62% at 5 years,71% versus 55% at 10 years) [12]. Nevertheless, vascular tissueengineering has become a promising approach to overcomethe limits of autografts (e.g.,morbidity and scarce availability)and the inappropriate properties of synthetic grafts.

This review summarizes the most relevant and recentstudies on vascular tissue engineering for small diameterblood vessel regeneration, focusing on the development ofscaffolds made of decellularized matrices and natural and/orbiodegradable synthetic polymers and on the realization oftissue-engineered vascular grafts (TEVGs) without scaffold(Figure 1).

2. Vascular Tissue Engineering

As reported by Couet et al. [6], vascular tissue engineering“aims to apply the principles of engineering and life sciencestowards the development of a vascular construction thatdemonstrates biological and mechanical properties as closeas possible to those of a native vessel”.

The requirements of an ideal tissue-engineered vasculargraft (TEVG), with both large or small diameter, are summa-rized in Table 2 [3, 5, 13, 14]. Among all requirements for anideal TEVG (Table 2), the strictest requisites are correlatedto the regeneration of a functional endothelium and thesimilarity between the mechanical proprieties of TEVG andnatural blood vessels. These two requirements are related tothe failure of PET and ePTFE prostheses for small calibervessels.

The basic strategy for vascular tissue engineering consistsof the design and the production of appropriate scaffolds forvascular cell adhesion, proliferation, and differentiation andthe choice of cell type. For human applications, the idealcells should be nonimmunogenic, functional, and easy toisolate and expand [10]. Two different approaches are mainlydeveloped; the first method consists of the bioreactor uses togenerate physiological-like stimuli onto cell seeded scaffoldsfor in vitroTEVGmaturation, before the in vivo implantation.The latter approach regards the direct implantation of cell

seeded scaffolds in the body that acts as a bioreactor forTEVG maturation. Recently, some studies are focused onthe need for off-of-the-shelf grafts for the regeneration ofsmall diameter blood vessels, analyzing the possibility todirectly implant acellular scaffolds in the body. The aim ofthis approach is to develop readily available grafts for urgentvascular surgery.

In the last years increasing interest has been paid todevelop an appropriate substitute for small diameter bloodvessel replacement. The next paragraphs are focused onthe most relevant and recent approaches for vascular tissueregeneration.

2.1. Scaffolds from Decellularized Matrices. Decellularizationprocess aims to remove all cellular and nuclear matter min-imizing any adverse effects on the composition, biologicalactivity, andmechanical integrity of the remaining extracellu-larmatrix (ECM) for the development of a new tissue [15–17].The process usually consists of mechanical shaking, chemicalsurfactant treatment, and enzymatic digestion [4]. As poten-tial sources of ECM, many organs and tissues (such as skin,ureter, and liver) from humans and animals (such as bovine,sheep, monkeys, pigs, and rabbits) have been decellularizedfor different applications, such as skin, bone, and valvularheart regeneration [15, 16]. Decellularized matrix advantagesare correlated to its natural three-dimensional ultrastructureand its structural and functional proteins, essential for celladhesion, migration, proliferation, and differentiation [10,17, 18]. However, the specific composition of the ECM isrelated to the tissue source. ECM matrices demonstrated tobemainly affected by the age and health status of the animal atharvest and by the manufacturing process, influencing theirquality, mechanical and biochemical properties, biocompati-bility, and clinical performance [15].The presence of potentialantigenic compound traces (e.g., lipids, DNA, and glycosy-lation products) may cause an inflammatory response [15,17]. Furthermore, decellularization procedures may removedesirable ECMcomponents, such as collagen, thus decreasingmechanical properties [15–17].Decellularizedmatrices can bestored in hydrated state or in dehydrated lyophilized form.Hydrated ECM matrices demonstrate excellent biomechan-ical characteristics and improved cellular ingrowth rates; incontrast lyophilized ECM matrices show long shelf life andeasy transportability [15, 17].

From a historical context, decellularizedmatrices derivedfrom many animal organs and tissues (such as porcinesmall intestinal submucosa, porcine aortas) were widelyinvestigated. In 1999, Sullivan’s group [19] developed TEVGsmade of porcine small intestinal submucosa and type I bovinecollagen. Specifically, porcine small intestinal submucosa waschemically decellularized, wrapped around a 4mmmandrel,and impregnated with bovine collagen in the TEVG lumen[19].Then, the collagen layer was crosslinked and coated withheparin-benzalkonium chloride complex [19]. TEVGs wereinterpositionally implanted in the common carotid artery ofrabbits for 90 days [19].These TEVGs demonstrated excellentpatency without hyperplasia and aneurysm formations andallowed the SMC cellularization and the endothelialization,showing physiological vasoreactivity to agonists [19]. In 2000,

ISRN Vascular Medicine 3

Table 1: Compliance of natural vessels and synthetic prostheses.

Graft type Compliance Elastic modulus Reference

Artery 5.9 ± 0.5%/100mmHg — [41]677%/MPa 0.455MPa [42]

Saphenous vein 4.4 ± 0.8%/100mmHg — [41]

PET 1.9 ± 0.3%/100mmHg — [41]145%/MPa 1.9MPa [42]

ePTFE 1.6 ± 0.2%/100mmHg — [41]124%/MPa 2.2MPa [42]

Vascular tissue engineering

Decellularizedmatrices

Synthetic polymers

PGA, PLA, PCL, and PGS

Hybrid scaffoldsCompletely

Fibrin, elastin,hyaluronan, silk In vivo bioreactor,

cell sheet-based,and cell ring-based

approaches

Naturalpolymers biological

TEVGs

fibroin, andcollagen

Figure 1: Sketch of the review content.

Table 2: Requirements of an ideal TEVG, in particular small diam-eter vessels.

Biocompatibility

NontoxicityNonimmunogenicityNonthrombogenicityNonsusceptibility to infectionAbility to grow for pediatric patientsMaintenance of a functional endothelium

Mechanical properties

Compliance similar to native vesselBurst pressure similar to native vesselKink and compression resistanceGood suture retention

Processability

Low manufacturing costsReadily available with a large variety oflengths and diametersSterilizableEasy storage

Haverich’s group [20] developed a trypsin-based decellular-ization procedure to remove cells from porcine aortas and,then, a recellularization method using a bioreactor, humanperipheral venous ECs, and myofibroblasts. To study thein vivo immune response, decellularized aortas were sub-cutaneously implanted in the rat model showing a reducedpresence of t-lymphocytes and leukocytes in comparisonwith the control (not decellularized porcine aortas) [20].They reported for the first time the complete in vitro

endothelialization and the in vitro intramural myofibroblastrepopulation into decellularized matrices using a bioreactorand human cells [20]. In 2001, Mayer’s group [21] chemicallydecellularized porcine iliac vessels and noninvasively isolatedendothelial progenitor cells (EPCs) from peripheral blood ofsheep. EPCs were in vitro cultured and rotationally seededinto the decellularized matrices. After 4 days in a laminarflow bioreactor, decellularized vessels were in vivo implantedinto the common carotid arteries of sheep by an end-to-endanastomosis [21]. Decellularized vessels remained patent for130 days due to the presence of EPCs; in fact not seededdecellularized vessels occluded within 15 days [21]. Further-more, EPCs-decellularized matrices demonstrated contrac-tile activity and nitric oxidemediated vascular relaxation thatwas similar to native carotid arteries [21]. In 2003, for thefirst time, Niklason’s group [22] developed a decellularizationstrategy for TEVGs obtained by culturing bovine aortic orporcine carotid SMCs onto polyglycolic acid (PGA) meshes.This approach avoided the use of allogeneic or xenogeneictissue, eliminating the risk of viral disease transmission dueto the possible utilization of highly screened cells [22]. Thegroup improved this approach in the next years as reportedin the following (Table 3).

Table 3 reports the most relevant and recent (2008–2013)studies on decellularized matrices for vascular tissue engi-neering, focusing on the material source, decellularizationmethod, and mechanical and biological performances.

Among all the approaches reported in Table 3, particularattention is paid to an interesting strategy developed byNiklason’s group [23] (Figure 2).This approach consisted of a


Table3:Stud

ieso

fthe

scientificliterature

(2008–2013)o

ndecellu

lariz

edmatric

esforv

asculartiss

ueengineering.

Organ/m

aterial

Decellulariz

ation

metho

dDecellulariz

ation

results

IDMechanicalresults

Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

Bovine

jugu

lar

vein

Multistepdetergent

enzymaticprocess+

photoo

xidativ

elycrosslink

ing

Integrity

ofcollagen

fibrilsa

ndelastic

fibers

Patch

UTS∼6∗

MPa

significant

high

erthan

nativ

ebo

vine

veins(∼5∗

MPa)

Coatin

gwith

fibronectin,

gelatin

,and

collagenIV;

SEM

imagesandhisto

logical

analysis(7

days):

confl

uent

layer

ofHUVEC

s

Mod

el:rat(12weeks);

implantedsubcutaneously;

graft

stability;

chronicinfl

ammatory

respon

se

[59]

Rataortic

cond

uitg

raft

Detergent-based

protocol

——

——

Mod

el:rat(8weeks);

fibronectin-coatedgraft

s(on

both

graft

surfa

ces);

fibronectin

surfa

cecoating

persistentfor

atleast8

weeks;

luminalendo

thelialization

acceleratedby

fibronectin;

localm

yofib

roblast

hyperplasia

increasedby

fibronectin;

cellinvasio

nfro

mthe

adventitiallayerintothe

mediaincreasedby

fibronectin;

enhanced

matrix

metalloproteinase

activ

ityby

fibronectin;

noinflammatorycellmarkers;

nosig

nsof

thrombo

sis

[60]

Urin

arybladder

Soakingin

salin

e+mechanical

delamination+soaking

inph

osph

ate-bu

ffered

salin

e[61]

Intactbasement

mem

brane;

bimod

alsurfa

cePatch

1layer:

UTS

=1.9

–2.3MPa,

𝜀𝑏=38–4

0%;

4layers:

UTS

=21–30M

Pa,

𝜀𝑏=38–4

0%

SEM

andconfocalmicroscope

images(5

days):

excellent

adhesio

n,spread,

andproliferatio

nwith

phenotypep

reservationof

HAo

ECsa

ndHAo

SMCs

—[62]

Porcine

abdo

minalaorta

andcarotid

Detergent-enzym

atic

process

PreservedEC

Mstructurew

ithno

resid

ualcells;

removalof

the

majority

ofDNA

content

2–11mm

Circum

ferentia

ltensile

tests

𝐸=0.22

aMPa,

UTS

=2.01

aMPa,

𝜀𝑏=1.3

5a,

𝐶=2.27×10−3a

1/mmHg,

BP=2560

ammHg,

SRS=732ag;

——

[18]

ISRN Vascular Medicine 5Ta

ble3:Con

tinued.

Organ/m

aterial

Decellulariz

ation

metho

dDecellulariz

ation

results

IDMechanicalresults

Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

stress-r

elaxatio

ntests

𝜎resid

ual=0.78%;

similartona

tivep

orcine

arteries

(except𝜀𝑏,

𝜎resid

ual):

𝐸=0.19

aMPa,

UTS

=1.5

5aMPa,

𝜀𝑏=1.8

3a,

𝐶=2.61×10−3a

1/mmHg,

BP=2331

ammHg,

SRS=882akP

a,𝜎resid

ual=0.58%

Porcine

abdo

minalaorta

Non

ionicd

etergent

process+

lyop

hilization

Intactcollagenand

elastin

7-8m

m—

Smoo

thmuscle

-like

and

endo

thelial-likec

ells

differentiatedin

vitro

from

autologous

BMMNCs

(3weeks)

Mod

el:pig(18weeks);

before

implant:seedingof

SMCs

andEC

songraft

s(1

weekof

cultu

re);

nosig

nof

thrombu

sform

ation,

dilatation,

orste

nosis

;regeneratio

nof

endo

thelium,

tunica

media,and

adventitia;

presence

ofcollagenand

elastin

;grow

thpo

tential

[63]

Hum

anum

bilicalartery

Detergent

process

Integrity

ofextracellular

collageno

usmatrix

;maybe

partial

redu

ctionof

elastin

1.5mm

BP=840.37±

114.67m

mHg,

UTS

=1618.21±

691.2

6kPa,

𝐸=7.4

1±3.85

MPa,

𝐶=4.26±

2.96%/10

0mmHg;

similartonativeh

uman

umbilicalarterie

s:BP

=969.6

6±

154.42

mmHg,

UTS

=1372.23±

809.3

0kPa,

𝐸=13.33±6.85

MPa,

𝐶=5.84±

3.10%/10

0mmHg

—

Mod

el:rat(8weeks);

acellularg

rafts

implanted;

5ratsdied

with

infewho

urs

after

implantatio

ndu

eto

thrombo

sis;

other6

TEVG

srem

ained

patent;

thrombo

sisatthep

roximal

anastomosis;

noruptureo

raneurysm

form

ation

[64]


Table3:Con

tinued.

Organ/m

aterial

Decellulariz

ation

metho

dDecellulariz

ation

results

IDMechanicalresults

Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

PGAtubu

lar

scaffoldcultu

red

with

human

allogeneicor

canine

SMCs

usinga

bioreactor

(7–10

weeks)

Detergent

process

—

3,4m

mcanine

TEVG

6mm

human

TEVG

Canine

TEVG

sBP

=1618±67

bmmHg;

human

TEVG

sSR

S=178±11

bg;

BP=3337±343b

mmHg;

C=3.3±

0.8b%/10

0mmHg;

similartonativ

ehum

anbloo

dvessels

Canine

TEVG

sbefore

implantatio

n:cultu

rewith

autologous

ECs;

∼21

days

forE

Cexpansion

and2days

forseeding

and

precon

ditio

ning

ina

bioreactor

Canine

TEVG

sfor

coronary

artery

bypass

mod

el:dog

(12mon

ths);

excellent

long

-term

patency,

andno

steno

sisor

dilatatio

n,andno

intim

alhyperplasia

;1animaldied

with

apatent

graft

;1TE

VGocclu

dedat1

week;

canine

TEVG

sfor

perip

heral

artery

bypassmod

el:dog

(1mon

th);

only1animaldied

with

apatent

TEVG

(1day);

human

TEVG

sfor

arterio

venous

accessfor

hemodialys

is;mod

el:babo

on(6

mon

ths);

patency:88%(7

of8);1

TEVG

with

thrombo

sisat3mon

ths;

noaneurysm

aldilatatio

n,and

nocalcificatio

n,andno

substantialintim

alhyperplasia

[23]

UTS

:ultimatetensilestre

ngth;𝜀𝑏:strainatbreak;𝐸:elasticm

odulus;B

P:bu

rstp

ressure;𝐶:com

pliance;SR

S:sutureretentionstr

ength;∗

values

wereg

raph

icallyread;amedian;

b mean±sta

ndarderroro

fthe

mean;

𝜎:stre

ss.

HUVEC

s:hu

man

umbilicalvein

endo

thelialcells;

HAo

SMCs

:hum

anaorticsm

ooth

muscle

cells;H

AoEC

s:hu

man

aorticendo

thelialcells;

BMMNCs

:bon

emarrowmon

onuclear

cells.


first step for growing allogeneic SMCs onto a PGA tubularscaffold in a bioreactor (Figure 2(a)). The PGA scaffoldrapidly degrades, while cells secrete ECM proteins for tissueregeneration (Figure 2(b)). After the maturation period (7–10weeks), the construct was decellularized and stored in a buffersolution at 4∘C until the in vivo implantation (Figure 2(c)).TEVGs with ID ≥ 6mmwere readily available for the implant(Figure 2(d)); in contrast TEVGs with ID = 3-4mm wereseededwith autologous ECs isolated by a biopsy (Figure 2(e)),before implantation, to provide an antithrombogenic luminalsurface,minimizing the risk of graft occlusion [23].Niklason’sgroup used PGA scaffolds only as a support for the ECMdeposition; therefore, the final decellularized constructs didnot include the synthetic polymer. This approach may allowthe production of many TEVGs using only one humandonor [23]. In addition, TEVGs may be produced withthe appropriate diameter and ready available due to use ofallogeneic cells and decellularization [23].

2.2. Scaffolds from Natural Polymers. Natural polymers gen-erally show excellent biological performances; specifically,they do not activate chronic inflammation or toxicity [24].

Among the natural polymers studied for application invascular tissue engineering, this paragraph is focused onfibrin, elastin, hyaluronan, silk fibroin, and collagen thatshow interesting properties for vascular tissue engineeringapplications and are now the main studied natural polymers.

Fibrin is an insoluble body protein entailed in woundhealing and tissue repair [6, 25]. Fibrin clot, obtained byfibrinogen polymerization due to thrombin, is a fibrillarnetwork gel that provides a structural support for adhesion,proliferation, and migration of cells involved in the healing[6, 25]. Finally, fibrin clot is resorbed through the fibrinolysis,a fibrinolytic process that breaks down fibrin fibrils [6, 25].Fibrinogen may be purified from autologous blood and usedfor scaffold fabrication avoiding immunological problems[6, 25].

Elastin is one of the major ECM proteins in the arterialwall [6, 25–27] that confers elastic recoil, resilience, anddurability [26–28]. It is an important autocrine regulator toSMC and EC activity, inhibiting migration and proliferationof SMCs and enhancing attachment and proliferation of ECs[26, 28]. Elastin, as a coating of vascular devices (made ofePTFE [29], PET, a copolymer of ePTFE and polyethylene,and a polycarbonate polyurethane [30]), demonstrated lowthrombogenicity with reduced platelet adhesion and activa-tion [26, 27].

Hyaluronan is an anionic nonsulfated glycosaminogly-can (GAG) that consists of glucuronic acid and N-acetylglucosamine [6]. It can be produced in large amount bymicrobial fermentation [25]. Furthermore, hyaluronic acid ishydrophilic, nonadhesive, biocompatible, and biodegradable[6, 25].

Silk fibroin is a protein produced by silkworms andspiders [31]. The amino acid structure of silk fibroin fromBombyx mori is composed mainly of glycine (43%), alanine(30%), and serine (12%) [31]. It shows excellent mechanicalproperties and biocompatibility [31]. Silk degrades slowly and

(a)

(b)

(c)

(d)

(e)

Figure 2: Sketch of the innovative strategy developed by Niklason’sgroup, from [23] (reprinted with permission from AAAS).

proteolytically in vivo, maintaining more than 50% of itsmechanical properties after 2 months [32].

Collagen is the major ECM protein in the body thatsupplies mechanical support to many tissues [6]. Collagendemonstrates low antigenicity, low inflammatory response,biocompatibility, biodegradability, and excellent biologicalproperties [6, 25, 33]. Collagen type I is one of the maincomponents of the vascular wall, whereas it is widely used asscaffold for vascular tissue engineering applications [6].

Seminal studies on the use of natural polymers for TEVGare concerned with the development of scaffolds made ofcollagen or fibrin. In 1986, Weinberg and Bell reported thedesign and the fabrication of the first TEVGmade of collagengel and cells [34]. Specifically, they developed a tubular graftby casting and gelling bovine aortic SMCs and collagen inan annular mold at 37∘C, to mimic the tunica media [34].Then, a mixture of collagen and bovine aortic adventitialfibroblasts was cast around the first tube to recreate thetunica adventitia and, finally, bovine aortic ECs were seededinto the lumen [34]. Unfortunately, the burst pressure of theobtained TEVG was very low and less than 10mmHg [34].Therefore, a PET mesh was added allowing for an increaseof the burst pressure (40–70mmHg) [34]. Furthermore, aTEVG composed of three layers of collagen gel and two PETmeshes reached a burst strength of 120–180mmHg [34]. In2000, Nerem’s group [35] designed a dynamic mechanicalconditioning method to improve the mechanical propertiesof adult rat aortic SMCs entrapped in a collagen-gel scaffold.The TEVG was cultured onto an inflatable silicone tubeallowing for the transmission of the mechanical stimuli (i.e.,


cyclic strain) [35]. Due to the mechanical conditioning,SMCs increased their circumferential orientation, leadingto an improvement of mechanical properties (higher yieldstress, ultimate stress, and elastic modulus) [35]. In 2002,Tranquillo’s group [36] demonstrated that neonatal aortic ratSMCs embedded in fibrin gel were stimulated to increasethe collagen production in comparison with SMCs entrappedin collagen gel. Furthermore, they inhibited the rapid fibrindegradation by SMCs, due to the addition of 𝜀-aminocaproicacid to the culture medium that avoids the binding ofplasmin or plasminogen to fibrin [36]. After 6 weeks of invitro incubation in TGF-𝛽 and insulin, tubular fibrin gelwith entrapped SMCs exhibited ultimate tensile strength andelastic modulus similar to those of rat abdominal aorta [37].In 2005, Andreadis’s group [38] developed scaffolds basedon ovine SMCs embedded in fibrin gels, using aprotinin asfibrinolysis inhibitor. After 2 weeks in culture, ovine ECswereseeded on the outer surface of the TEVG and cultured for3 or 10 days [38]. After the reversion of TEVGs (i.e., ECsin the TEVG lumen), TEVGs were interpositionally in vivoimplanted in jugular veins of 12-week-old lambs for 15 weeks[38]. TEVGs remained patent and showed blood flow ratessimilar to those of the native jugular vein [38]. Furthermore,TEVGs allowed the production of elastin and collagen fibers,the SMC circumferential alignment, and the presence of auniform endothelium [38].

In the last years (2008–2013, Table 4), many researchgroups studied the possible use of natural polymers forTEVGs, following the first results previously obtained [34–38].

Among the number of recent studies developed to fab-ricate TEVGs using natural polymers, particular attention ispaid to structural proteins, such as fibrin [39, 40], elastin[27], and collagen [33], and to the methods to improve theirmechanical properties [39, 40] (Table 4). Simultaneously, silkfibroin is widely investigated for vascular application dueto its higher mechanical properties in comparison to othernatural polymers, such as fibrin [39] (Table 4). However, it isdifficult to compare mechanical properties among differentstudies because of the different conditions used for mechan-ical characterization, such as crosshead speed, sample shape(rectangular, tubular), and applied load direction.

2.3. Scaffolds from Biodegradable Synthetic Polymers. Biodeg-radable synthetic polymers generally demonstrate tailorablemechanical properties and high reproducibility and, com-pared to natural polymers, can be produced in large amounts[27, 43].

Among the biodegradable synthetic polymers understudy for application in vascular tissue engineering, this par-agraph is focused on polyglycolic acid (PGA), polylacticacid (PLA), poly-𝜀-caprolactone (PCL), and polyglycerol-sebacate (PGS).

PGA is a semicrystalline, thermoplastic aliphatic polyes-ter synthesized by the ring-opening polymerization of gly-colide [6, 25]. It degrades rapidly in vivo by hydrolysis toglycolic acid, metabolized and eliminated as carbon dioxideand water, and completely degrades in vivo within 6 months[6]. PGA is a Food andDrugAdministration (FDA) approvedpolymer [10] for human clinical use.

PLA is a thermoplastic aliphatic polyester synthesized byring-opening polymerization of lactic acid [6, 25]. It demon-strates good biocompatibility and mechanical properties andthe ability to be dissolved in common solvents for processing[44]. PLA ismore hydrophobic than PGA, leading to a slowerdegradation rate [6, 25]. It is a FDA approved polymer [10] forhuman clinical use. PLA is a chiral molecule: poly-D-lactide(PDLA) and poly-L-lactide (PLLA) are the enantiomericsemicrystalline forms; in contrast the racemate, poly-D,L-lactide (PLA), presents an amorphous structure [6, 44]. PLLAtakes months or even years to lose its mechanical integrity[25].

PCL is a semicrystalline, aliphatic polyester synthesizedby the ring-opening polymerization of 𝜀-caprolactone [6, 25,45]. It shows good mechanical properties, specifically highelongation and strength, and good biocompatibility [45, 46].Furthermore, PCL degrades very slowly in vivo by enzymaticaction and by hydrolysis to caproic acid and its oligomers[6, 45]. It takesmore than 1 year to completely degrade in vivo[45, 46]. PCL is a FDA approved polymer [6].

PGS is an elastomer synthesized by polycondensation ofglycerol and sebacic acid [47]. It demonstrates good biocom-patibility and good mechanical properties, specifically highelongation and low modulus, indicating an elastomeric andtough behavior [47]. PGS degrades in vivo by hydrolysis in2 months [47]. The FDA approved glycerol and polymerscontaining sebacic acid for medical applications [47].

Seminal studies are mainly concerned with the devel-opment of scaffolds made of PGA. In 1997, Langer’s group[49] designed PGAmesh tubular scaffolds coated with PLLAor 50/50 PLGA copolymer to stabilize PGA meshes. TEVGswere implanted into the omentum of rats demonstrating thein vivomaintenance of their structure. Furthermore, TEVGswere in vitro seeded and cultured with bovine aortic SMCsand ECs, demonstrating good adhesion and cell prolifera-tion up to 14 days. In 1999, they developed PGA scaffoldschemically modified with sodium hydroxide to increasehydrophilicity [50]. Bovine aortic SMCs were in vitro seededand cultured onto PGA scaffolds under pulsatile radial stressusing a bioreactor [50]. After 8 weeks, SMCs migrated in theTEVG wall thickness and bovine aortic ECs were seeded andcultured into the TEVG lumen under continuous perfusion[50]. The TEVG demonstrated good mechanical strength[50]. TEVG was interpositionally implanted into the rightsaphenous artery of miniature swine and remained patentfor 4 weeks, even though there was a decrease of the bloodflow [50]. Furthermore, histological analysis showed highlyorganized structurewithminimal inflammation [50]. In 1999,Mayer’s group [51] designed scaffolds made of a copolymerof PGA and polyhydroxyalkanoate (PHA). Specifically, ovinecarotid arteries were harvested and cultured in vitro [51].After 6–8 weeks, autologous mixed cell population of ECs,SMCs, and fibroblasts was seeded and cultured onto TEVGs[51].Then, TEVGs were implanted in the abdominal aortas oflambs for 5 months [51]. All TEVGs remained patent withoutaneurysm formations and histological analysis showed thepresence of elastic fibers in the tunica media and ECs inthe tunica intima [51]. During the implantation period, theTEVG mechanical properties changed and became similar


Table4:Stud

ieso

fthe


(2008–2013)o

nTE

VGsfabric

ated

with

naturalp

olym

ers.

Material

Fabricationmetho

dID

Mechanicalproperties

Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

Fibrin

Two-layerg

rafts:SMCs

embedd

edin

fibrin

hydrogel+fib

rinlay

ercompo

sedof

high

concentrationfib

rinogen

4mm

After2we

eks

twolayers

UTS∼90∗

kPa,

BP=177±5.3m

mHg;

singlelayer

UTS∼65∗

kPa,

BP=18±7.7

mmHg;

UTS

calculated

from

circum

ferentialtensiletests

SMCs

isolatedfro

m3-day-old

lambs;

histo

logicalanalys

isaft

er2weeks

ofcultu

re:

SMCs

distrib

uted

unifo

rmlyin

theinn

erlay

er,SMCs

didno

tmigrateinto

theo

uter

fibrin

layer,andSM

Csmaintainedtheir

contractile

prop

ertie

sinsideg

els

—[39]

Fibrin

nDFentrapmentin

fibrin

gel+

static

cultu

refor2

weeks

+dynamicbioreactor

cultu

refor5

–7weeks

(stim

uli:cyclic

stretching

+luminal,

ablumenal,transmural

flows)

2.4m

m

After7–9we

eks

circumfer

entia

ltensiletests

𝐸∼2500∗

kPa,

UTS∼1800∗

kPa;

axialtensiletests

𝐸∼900∗

kPa,

UTS∼1200∗

kPa;

ID=2m

m,

BP=1366±177m

mHg,

SRS=0.19±0.05

N,

𝐶∼4.5∗%/10

0mmHg;

ID=4m

mBP

=1542±188m

mHg,

SRS=1.3

2±0.58

N(w

ithentrappedPL

Acuffs),

𝐶∼2.5∗%/10

0mmHg;

ovinefem

oralartery

BP=2297±207m

mHg,

𝐶∼3.25∗

%/10

0mmHg

Histologica

lanalys

isaft

er7–9

weeks

ofsta

tic+dynamic

cultu

re:circ

umferentialcollagen

fiber

alignm

ent,

minor

resid

ualfi

brin

in2m

mgraft

s,no

neevidentin4m

mgraft

s,andnh

DFs

evenly

distrib

uted

acrossthethickness

—[40]

Silkfib

roin

Electro

spinning

usinga

rotatin

gtransla

ting

mandrel

6mm

BP=575.67±17.47

mmHg,

𝐶=3.51±0.42%/10

0mmHg;

circumfer

entia

ltensiletests

DR=5m

m/m

in:

UTS∼2.4∗

N,

𝜀𝑏∼57∗

%;

DR=25

mm/m

in:

UTS∼1.8∗

N,

𝜀𝑏∼37∗

%;

DR=50

mm/m

in:

UTS∼1.2∗

N,

𝜀𝑏∼34∗

%

MTT

assay(

7days):

linearincreaseo

fNIH

/3T3

viabilityover

time;

SEM

images(7

days):

NIH

/3T3

attached,spread,and

grew

;cellm

otilitythroug

hmatrix

—[65]

10 ISRN Vascular MedicineTa

ble4:Con

tinued.

Material

Fabricationmetho

dID


Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

Silkfib

roin

Gelspinning

1,1.5

mm

Silk

𝐸=2.20±0.90

MPa,

UTS

=0.273±0.11MPa;

nativerat

aorta

𝐸=2.44±0.76

MPa,

UTS

=0.519±0.11MPa;

PTFE

𝐸=918±52.9MPa,

UTS

=43.4±4.6M

Pa

Protein

adsorptio

n:lower

onsilkthan

PTFE

;platele

tadh

esionandspreading:

silkandPT

FErelativ

elyno

nthrom

bogenic;

confocalim

ages(overnight

incubatio

n):

high

erHCA

SMCandHUVEC

attachmento

nsilkthan

PTFE

;Pico

GreenDNA

assay(

for14

days):

high

errateof

HCA

SMCand

HUVEC

proliferatio

non

silk

than

PTFE

Mod

el:rat(4weeks);

noocclu

sion,

clotting

,or

ischemia;

vascular

cellremod

eling

:confl

uent

endo

thelium,

SMCs

migratio

nand

proliferatio

n

[66]

Silkfib

roin

+collagentype

I

Electro

spinning

usinga

rotatio

nalm

andrel+

dynamicdipp

ing

6mm

BP=894.00±24.91m

mHg,

𝐶=3.24±0.58%/10

0mmHg;

circumfer

entia

ltensiletests

DR=5m

m/m

inUTS

=3.17±0.32

N,

𝜀𝑏=58.96±4.07%;

DR=25

mm/m

inUTS

=2.76±0.39

N,

𝜀𝑏=35.6±2.13%;

DR=50

mm/m

inUTS

=2.5±0.12N,

𝜀𝑏=39.04±5.31%

MTT

assay(

7days):

linearincreaseo

fNIH

/3T3

viabilityover

time,

collagenincrease

metabolic

activ

ityatthee

arlytim

epoints;

SEM

images(7

days):

NIH

/3T3

attached,spread,and

grew

—[33]

Hyaff-11

(hyaluronan-

based

material)

Coatin

g+coagulation

bath

(ethanol)

4mm

——

Mod

el:porcine

(5mon

ths);

fibrin

glue:sealed

anastomoticsites;

anim

alsurvival:100%,n

osig

nsof

vascular

insufficiency;

2graft

scom

pletely

occlu

ded

after

2and3mon

ths,1g

raft

partially

occlu

dedaft

er4

mon

ths,and7graft

ssho

wed

noocclu

sionor

intim

alhyperplasia

;graft

salm

ostcom

pletely

degraded;

form

ationof

neotissue

(SMCs

,collagen,

elastin

fiberso

rganized

inlay

ers,

completee

ndotheliu

m)

[28]


Table4:Con

tinued.

Material

Fabricationmetho

dID


Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

Recombinant

human

tropo

elastin

#

Electro

spinning

usinga

rotatin

gtransla

ting

mandrel

4mm

1cm

𝐶=20.2±2.6%

/100m

mHg,

BP=485±25

mmHg;

circumfer

entia

ltensiletests

UTS

=0.34±0.14MPa,

𝜀𝑏=79±6%

,𝐸=0.15±0.05

MPa;

axialtensiletests

UTS

=0.38±0.05

MPa,

𝜀𝑏=75±5%

,𝐸=0.15±0.03

MPa;

nativ

eporcine

carotid

arteries

𝐶=3.4±0.5%

/100m

mHg;

circumfer

entia

ltensiletests

UTS

=2.59±0.31MPa,

𝜀𝑏=125±15%,

𝐸=0.41±0.09

MPa;

axialtensiletests

UTS

=0.95±0.13MPa,

𝜀𝑏=105±11%,

𝐸=0.20±0.06

MPa

Confocalim

agesaft

er48

hof

cultu

re:

BMEO

Csattached,spread,and

grew

;cellsform

edac

onflu

ent

mon

olayer

—[27]

# mon

omer

unitof

elastin

,cross-link

edtro

poelastin

mim

icnativ

eelastinfib

ers.

UTS

:ultimatetensiles

treng

th;𝜀𝑏:strainatbreak;𝐸:elasticm

odulus;𝐶

:com

pliance;BP

:burstpressure;SRS

:suturer

etentio

nstr

ength;DR:

deform

ationrate;∗values

wereg

raph

icallyread.

nDFs:neonatalhum

anderm

alfib

roblasts;

BMEO

Cs:porcine

bone

marrow-derived

endo

thelialoutgrow

thcells;N

IH/3T3

:murinefi

brob

lastcelllin

e;HUVEC

s:hu

man

umbilicalvein

endo

thelialcells;

HCA

SMCs

:hu

man

coronary

artery

smoo

thmuscle

cells.


to those of native vessels [51]. In 2001, Shin’oka’s group [52]performed the first transplantation of a TEVG in four-year-old girl. The TEVG consisted of PGA mesh coated with a50 : 50 copolymer of L-lactide and 𝜀-caprolactone, previouslyin vitro seeded with autologous vascular cells isolated fromperipheral vein biopsies [52]. After 7 months, there was noevidence of aneurysmor stenosis formations [52].The clinicaltrial performed by Shin’oka’s group is widely described inSection 4.

Among the several recent studies (2008–2013, Table 5)developed to fabricate TEVGs using biodegradable syntheticpolymers, particular attention is paid to PCL due to itsdeformability and mechanical strength (Table 5). PCL wasstudied alone [46, 53, 54], as multilayers (PGA-PCL-PGA[55]), electrospun coating on PGS [56], or copolymer withL-lactide [57, 58] (Table 5). For example, the use of PCLas PGS tube coating increased the mechanical properties ofthe substrate, specifically suture retention strength, elasticmodulus, and ultimate strength stress [56].

2.4. Synthetic versus Natural Scaffolds. TEVGs should haveappropriate biological behavior andmechanical properties toreach the clinical use. TEVG biological performance shouldsupport the complete integration of the grafts in the humanbody, avoiding the induction of a chronic inflammatoryresponse during the material degradation process. Further-more, TEVG mechanical properties should match those ofthe native blood vessels, specifically in terms of deformability,compliance, and strength.

One of the essential principles of tissue engineeringis the choice of a scaffold with appropriate degradationrate during the in vivo remodeling process. The speed ofthe scaffold degradation has to match the tissue-dependentregeneration rate to develop a functional substitute. Forvascular tissue engineering, the key point is the maintenanceof the structural and mechanical integrity of the neovasculartissue over all the regeneration process.

Generally, natural polymers lack mechanical perfor-mance, but they exhibit excellent biological behavior. Asreported in Table 4, tropoelastin electrospun scaffoldsshowed low mechanical strength compared to native bloodvessels, specifically, in terms of ultimate tensile strength(0.34 ± 0.14MPa) and burst pressure (485 ± 25mmHg)[27]. However, in vitro cell culture demonstrated a goodinteraction with porcine bone marrow-derived endothelialoutgrowth cells, forming a confluent monolayer after 2 days[27].

Normally, natural scaffolds are seeded and cultured invitrowith cells prior to in vivo implantation to obtain a partialremodeling of the structure due to ECM deposition. Thisin vitro remodeling usually leads to an increase of the graftmechanical properties. As reported in Table 4, fibrin gels withentrapped SMCs were dynamically cultured in a bioreactorfor several weeks to achieve a circumferential alignmentsimilar to native arteries [40]. This alignment improved themechanical properties of the grafts, reaching values similarto native blood vessels [40]. To date, studies related tonatural materials for vascular applications are focused on theimprovement of graft mechanical properties. To reach this

goal, some approaches consist of tissue in vitro maturation[40]. Other methods combine the use of natural materialswith synthetic polymers (such as tropoelastin with PCL [26]),reaching sometimes promising results in terms ofmechanicalproperties and cell/material interaction.

Synthetic polymers show controllable physical andmechanical properties but lack in biological performances.The products of PGA degradation could lead to a SMCdedifferentiation and to a pH decrease, potentially inducingan undesired inflammatory response [10, 39, 68]. Forexample, as reported in Table 5, PCL electrospun scaffoldsdemonstrated high elongation, similar to that of nativearteries, and slow degradation, allowing for cell infiltrationand tissue regeneration [46, 53, 54]. A long-term in vivostudy of the PCL electrospun tubes was performed in therat aorta for 18 months [46]. During the first 6 months, PCLelectrospun scaffolds induced a promising tissue regenerationwith a rapid colonization and a wide neovascularization[46]. In fact, fibroblasts and macrophages infiltrated andproliferated in the PCL graftmainly from the adventitial side;the fibroblasts become myofibroblasts that deposed a naturalECM in the PCL graft wall and the macrophages secretedangiogenic factors to induce neovascularization [46]. After6 months, tissue regeneration showed a clear regressionwith a reduction of capillaries and fibroblasts, an increaseof calcification, and a cell regression in intimal hyperplasia[46]. The tissue regression was caused by macrophagedisappearance, inducing the neovascularization decreaseand the consequent myofibroblast reduction, due to thedecrease of oxygen and nutrient presence [46]. Furthermore,calcifications were related to the process of chondroidmetaplasia that initially regards the transdifferentiationof smooth muscle cells into chondrocytes in the intimalhyperplasia layers [46]. Therefore, PCL electrospuntubes developed by de Valence et al. [46] demonstratedinsufficient regeneration of the blood vessels on the longterm (18 months); this result was unexpected because ofthe promising conclusions obtained in a previous study[79] regarding the implantation of PCL electrospun graft inrat model for 6 months. Specifically, the PCL electrospungrafts demonstrated patency, structural integrity, fasterendothelialization, ECM formation, and fiber degradationduring 6 months [79].

2.5. Hybrid Scaffolds from Synthetic and Natural Polymers.Recent studies have proposed hybrid scaffolds, made ofsynthetic and natural polymers, to combine the adequatemechanical properties of synthetic materials with the excel-lent biological behavior of natural polymers.

Table 6 reports the most relevant and recent (2008–2013)studies on TEVGs fabricated using hybrid scaffolds madeof natural and biodegradable synthetic materials, focusingon the fabrication method and mechanical and biologicalperformances.

Among the several recent studies aimed to fabricateTEVGs using hybrid scaffolds, particular attention is paid tothe use of PCL with natural polymers such as elastin [26, 68]and collagen [45, 68, 70] (Table 6). As an example, the addi-tion of PCL to tropoelastin electrospun scaffolds improved


Table5:Stud

ieso

fthe


(2008–2013)o

nTE

VGsfabric

ated

with

biod

egradables

yntheticpo

lymers.

Material

Fabricationmetho

dID


Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

PCL

Electro

spinning

of2

layersw

ithdifferent

porosityusinga

cylin

dricalrotatin

gtransla

tingmandrel;

ILLP

=innerlayer

with

lower

porosity;

OLL

P=ou

terlayer

with

lower

porosity

2mm

ILLP

UTS

=3.67±0.76

MPa,

𝜀𝑏=900±229%

,𝐸=8.8±0.9M

Pa,

BP=2845±241m

mHg,

SRS=556±47

gf,

WL=33.2±4.6m

Lcm−2min−1 ,

BL=0.25±0.06

mLcm−2min−1 ;

OLL

PUTS

=3.66±0.30

MPa,

𝜀𝑏=862±142%

,𝐸=8.5±0.8M

Pa,

BP=2688±192m

mHg,SR

S=490±41

gf,

WL=39.3±2.8m

Lcm−2min−1 ,

BL=0.89±0.39

mLcm−2min−1 ;

UTS

,𝐸,and𝜀𝑏calculated

from

axialtensiletests

—

Mod

el:rat(12weeks);

endo

thelializationnearly

complete;

ILLP

:allo

wed

good

cell

invasio

nfro

mthea

dventitia;

OLL

P:inhibitsgraft

cellu

lariz

ation;

perfe

ctpatencywith

nothrombo

sis

[53]

PCL

Electro

spinning

usinga

spinning

coun

ter

electrode

0.7m

m—

—

Mod

el:rat(18

mon

ths);

patency=72.5%;

noaneurysm

form

ation;

nocalcificatio

n;scaffolds

didno

tdisa

ppear

completely

;endo

thelium:com

plete

regeneratio

n;media:partia

lregeneration

[54]

PCL

Electro

spinning

usinga

cylin

dricalrotatin

gtransla

tingmandrel

2mm

BP=3280±280m

mHg,

SRS=936±32

g,WL=32.1±1.3

mLcm−2min−1 ,

BL=0.87±0.08

mLcm−2min−1 ;

axialtensiletests:

UTS

=4.1±

0.5M

Pa,

𝜀𝑏=1092±28%

—

Mod

el:rat(18

mon

ths);

graft

invivo𝐶=7.8±

0.9%

/mmHg,

proxim

alnativ

earteryin

vivo

𝐶=21.0±2.8%

/mmHg;

excellent

structuralintegrity:

noaneurysm

aldilation,

perfe

ctpatencywith

nothrombo

sis,and

limited

intim

alhyperplasia

(maxim

um5%

lumen

loss);

prom

ising

tissuer

egeneration

until

6mon

ths,aft

er6mon

ths

aclear

regressio

n;calcificatio

ns:spread

transm

urallyat18

mon

ths;

limitedvascularizationat18

mon

ths

[46]

14 ISRN Vascular MedicineTa

ble5:Con

tinued.

Material

Fabricationmetho

dID


Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

PGAor

PLLA

nonw

oven

felts

+PC

L/PL

Asealant

solutio

n

Dualcylinderc

hamber

molding

syste

m

PGA-

PCL/PL

A=0.9m

m;

PLLA

-PC

L/PL

A=0.7m

m

PGA-PC

L/PL

A:BP

=2710±282m

mHg,

SRS=3.13±0.72

N,

𝐸=33.0±7.0

MPa,

UTS

=5.3±0.72

MPa;

PLLA

-PCL

/PLA

:BP

=2790±180m

mHg,

SRS=4.37±0.67

N,

𝐸=24.0±5.9M

Pa,

UTS

=3.7±0.53

MPa

MTT

after

24h:no

differencein

HAo

SMCviabilitybetween

PGA-

PCL/PL

A,

PLLA

-PCL

/PLA

,and

TCP;

histo

logicalanalys

isaft

er72

h:HAo

SMCs

penetrated

and

adheredto

both

scaffolds

Mod

el:m

ice(6weeks);

acellularg

rafts

implanted;

foreignbo

dyreactio

nwith

giantcellformation(at3

weeks);

encapsulationarou

ndthe

externalperip

hery

ofPL

LA-PCL

/PLA

graft

s;no

thrombo

embo

liccomplications

oraneurysm

form

ation;

organizedvascular

neotiss

ue(end

otheliu

m,m

edia,and

collagen)

[57]

PGAno

nwoven

mesh+PC

Lpo

rous

sheet+

PGAno

nwoven

mesh

PGAandPC

Lseeded

with

SMCs

+aft

er4days

SMC-

PGAwou

ndarou

ndas

ilicone

tube

+aft

er3days

SMC-

PCL

wou

ndarou

ndSM

C-PG

A+aft

er3

weeks

PGAseeded

with

fibroblasts+aft

er2days

fibroblast-P

GAwou

ndarou

ndSM

C-PC

L(2

days)+

silicon

etub

eremoved

+seedingof

ECsinthelum

en+static

cultu

re(2

days)+

pulsa

tileb

ioreactor(14

days)

6mm

Circum

ferentia

ltensiletests

after

14da

ysof

cultu

re:

UTS

=827±155k

Pa,

𝐸of

collagenregion

=3.75±0.78

MPa,

𝐸of

elastin

region∼0.1∗MPa,

𝜀𝑏∼0.42∗

;on

lypo

lymericscaff

olds:

UTS

=91±21kP

a,𝜀𝑏∼0.6∗

;bo

vine

nativ

earteries:

UTS

=882±133k

Pa,

𝐸in

collagenregion

=3.31±0.56

MPa,

𝐸in

elastin

region∼0.075∗

MPa,

𝜀𝑏∼0.6∗

SMCs

,fibrob

lasts

,and

ECs

isolatedfro

mcalf;

histo

logicalanalys

is(afte

r14days

intheb

ioreactor):

form

ationof

3layers,

luminalarea:presenceo

fECs

,middlelayer:presenceo

fabun

dant

elastin

,well-organized

densec

ollagenEC

M,SMCs

,and

PCL,

outerregion:

presence

ofloosec

ollagenEC

M

—[55]

Heparin-coated

PGSpo

rous

tube

wrapp

edwith

aPC

Lthin

electro

spun

layer

Saltfusio

nmetho

d+

electro

spinning

720𝜇

m

PGS+PC

LSR

S=0.45±0.031N

,𝐸=536±119

kPa,

UTS

=3790±1450

kPa;

PGS

SRS=0.11±0.0087

N,

𝐸=243±71.8kP

a,UTS

=76.6±15.7kP

a;𝐸,U

TScalculated

from

circum

ferentialtensile

tests

;neoartery

BP=2360±673ammHg,

𝐶=11±2.2a%/10

0mmHg;

nativea

orta

—

Mod

el:rat(3

mon

ths);

noaneurysm

andste

nosis;

regu

lar,str

ong,and

synchron

ousp

ulsatio

nof

graft

s;confl

uent

endo

thelium,

contractile

smoo

thmuscle

layers,andcoexpressio

nof

elastin

,collagen,

andGAG

;

[56]


Table5:Con

tinued.

Material

Fabricationmetho

dID


Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

BP=3415±529a

mmHg,

𝐶=6.7±2.3a%/10

0mmHg

graft

extensived

egradatio

n+

ECM

synthesiz

ationwith

in14

days;

remod

eling

stillactiv

eat3

mon

ths

PCL/PL

A

Patie

nt-specific

scaffolds:

rapidprototyping(usin

gradiograph

icX-

ray

slices)+lostwax

+dip

coating+selective

dissolution+salt

leaching

>1m

m𝐸=0.6to

5.2M

Pa;higher𝐸

forlow

erpo

rosity

andhigh

erconcentration

PCL/PL

Ascaffolds

coated

with

collagen;

fluorescent

images(4

days):

HUVEC

attachment;

materialbiocompatib

ility

confi

rmed

—[58]

PGAno

nwoven

mesh

Seedingof

SMCs

+cultu

refor7

days

instaticcond

ition

+mesh

wou

ndarou

ndas

ilicone

tube

+pu

lsatile

bioreactor

(8we

eks)

4mm

After8we

ekso

fculture

UTS

=0.62×10

6Pa,

𝐸=10.5±1.2

5MPa,

SRS=1.2

6±0.16N,

BP∼1.2∗

MPa;

∼50–60%

ofthoseo

fhum

ansaphenousv

eins:

UTS∼0.95×10

6∗Pa,

𝐸∼14.5∗

MPa,

SRS∼2.25∗

N,

BP∼1.2∗

MPa;

UTS

,𝐸calculated

from

circum

ferentialtensile

tests

SMCs

differentiatedfro

mhA

SC(7

days);

histo

logicalanalys

is(afte

r8we

eksintheb

ioreactor):

noremnant

ofun

degraded

PGA

fibers,

denses

tructure

with

well-p

opulated

SMCs

,SM

Csorientated

inlay

ers,and

noelastic

fibers

—[67]

PCL/PL

A:50:

50copo

lymer

of𝜀-caprolacton

eand

L-lactide.

UTS

:ultimatetensile

strength;𝐸:elasticmod

ulus;𝜀𝑏:strainatbreak;SR

S:suture

retentionstr

ength;

BP:burstpressure;𝐶

:com

pliance;WL:

water

leakageat120m

mHg;BL

:blood

leakageat120m

mHg;∗

values

wereg

raph

icallyread;amean±sta

ndarderroro

fthe

mean.

hASC

s:hu

man

adipose-deriv

edste

mcells;H

AoSM

Cs:hum

anaorticsm

ooth

musclec

ells;

HUVEC

s:hu

man

umbilicalvein

endo

thelialcells;

GAG

:glycosaminoglycan;

TCP:

tissuec

ulture

plastic

.


Table6:Stud

ieso

fthe


(2008–2013)o

nTE

VGsfabric

ated

with

naturaland

synthetic

polymers.

Material

Fabricationmetho

dID


Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

Blendof

PCL+

elastin

+collagentype

I

Electro

spinning

of3layersatdifferent

polymer

ratio

susin

ga

cylin

dricalrotatin

gtransla

tingmandrel

2mm

SRS=91–189∘

gf,

BP=2387–>

3000∘

mmHg;

𝐶=2.83–0

.77∘%/10

0mmHg,

increase

ofPC

L+decrease

ofelastin

:increase

ofSR

SandBP

,decreaseo

f𝐶

——

[68]

Recombinant

human

tropo

elastin

(innerlayer)+

PCL(outer

layer)

Electro

spinning

usinga

rotatin

gmandrel;

cross-lin

king

:glutaraldehyde

vapo

rs

2.8m

m

𝐸∼300∗

kPa,

BP∼1900∗

mmHg,

𝐶∼0.065∗

kPa−

1 ,HP=minim

al;

similartointernalmam

marya

rtery:

𝐸∼267±46

kPa,

BP∼2267±215m

mHg,𝐶∼0.09∗

kPa−

1

SEM

andflu

orescencem

icroscope

(3days):

HUVEC

satta

ched

and

proliferated;

platele

tatta

chment:

redu

ced,comparedto

ePTF

Eand

PCL

Mod

el:rabbit(1m

onth);

maintenance

ofph

ysical

integrity

;no

evidence

ofdilatation,

anastomoticdehiscence,or

seroma;

mechanicalpropertieso

fim

plantedgraft

srem

ained

similartocontrols

[26]

PGA+PL

LA+

collagentype

I

Collagenmicrospon

ge+

woven

tube

madeo

fcore-sheathyarns(PL

LAandPG

A,resp.)

fabricated

byair-jet

spinning

4mm

Axialtensiletests

before

implan

tatio

n:UTS∼30∗

MPa

versus,

𝐸∼210∗

MPa;

after

implan

tatio

n:UTS∼5–9∗

MPa,

𝐸∼70–100∗

MPa;

nativ

ecarotid

artery:

UTS∼7.5∗

MPa,

𝐸∼60∗

MPa

Cellculture

with

NIH

/3T3

orHUVEC

(3days);

SEM

analysis:

high

ercellnu

mber

andbette

radh

esionforg

rafts

with

collagen

Mod

el:dog

(12mon

ths);

acellularg

rafts;

100%

patency,no

thrombo

sisor

aneurysm

;excellent

tissuer

egeneration:

completee

ndothelialization,

SMCs

,elastin,and

collagen

fibers;

PGAfib

ersc

ompletely

absorbed

at2mon

ths,PL

LAfib

ersu

nabsorbedat12

mon

ths

[69]

Blendof

PCL+

collagentype

IElectro

spinning

usinga

rotatin

gmandrel

4.75

mm

SRS=3.0±1.1

N,

BP=4915±155m

mHg,𝐶=5.6±

1.6%/10

0mmHg;

axialtensiletests:

UTS

=4.0±0.4M

Pa,

𝐸=2.7±1.2

MPa,

𝜀𝑏=140±13%;

maintenance

oftensile

prop

ertie

sinap

erfusio

nbioreactor

syste

mun

dero

verphysio

logical

cond

ition

sfor

4weeks;

nativ

eporcine

corona

ryartery:

UTS∼2.5∗

MPa,

𝐸∼1∗

MPa,

𝜀𝑏∼100∗

%

ECsa

ndSM

Csfro

mbo

vine

carotid

artery;

cytotoxicityv

iadirectcontact(7

days—MTS

assay):

nodifferences

incellviability

comparedto

TCPs;

cytocompatib

ility(7

days—MTS

assay):

ECandSM

Cadhesio

nand

proliferatio

n;cellcultu

reon

tubularg

rafts

(48h

—SE

Mandhisto

logical

analysis):SMCs

form

edmultip

lecelllayerson

theo

uter

region

,EC

sformed

amon

olayer

onthe

innersurface

—[45]


Table6:Con

tinued.

Material

Fabricationmetho

dID


Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

Blendof

PCL+

collagentype

IElectro

spinning

usinga

rotatin

gmandrel

4.75

mm

Circum

ferentia

ltensiletests

before

implan

tatio

n:UTS∼1.8∗

MPa;

after

implan

tatio

n:UTS∼0.8∗

MPa;

compa

rabletona

tivea

orta:

UTS∼1.2∗

MPa

ECsd

erived

from

endo

thelial

progenito

rcellsiso

latedfro

mlive

sheep;SM

Csiso

latedfro

msheep

artery

explants;

dynamiccultu

reusingap

ulsatile

bioreactor

underp

hysio

logic

cond

ition

s(9days);

SEM

+histo

logicalana

lysis:

ECsformed

acon

fluentlayer

ontheinn

erlumen,SMCs

assumed

amultilayered

confi

guratio

non

thee

xterior

Platele

tadh

esionin

alive

sheepmod

elfor15m

in:n

oadhesio

nforseededgraft

s,abun

dant

adhesio

nfor

unseeded

graft

s;model:

rabbit(1mon

th);

acellularg

rafts;

noaneurysm

aldegeneratio

n;majority

oftheg

rafts

remainedpatent

at1m

onth;

absenceo

finfl

ammatory

infiltrate

[70]

PCL/PL

A+

collagentype

I

PCL/PL

Aelectro

spinning

usinga

rotatin

gmandrel+

collagencoatingaft

erair

plasmatreatment

1,3m

m

Axialtensiletests:

UTS

=7.0±0.4M

Pa,

𝐸=16.0±7.1

MPa,

𝜀𝑏=289±55%;

circumfer

entia

ltensiletests:

UTS

=3.9±0.3M

Pa,

𝐸=16.6±4.4M

Pa,

𝜀𝑏=292±87%

HCA

ECsrotationally

seeded

intheb

ioreactorfor

4h+sta

ticcultu

re(10days—SE

Mand

histo

logicalanalysis

):HCA

ECse

venlydistrib

uted

and

wellspreadon

thelum

en

Mod

el:rabbit(7weeks);

acellularg

rafts

with

out

collagencoating;

nobloo

dleaking;

graft

structure

integrity

and

patency;

foreignbo

dyrespon

se:nocell

infiltration

[71]

PLDLA

+fib

rin

PLDLA

tubu

larw

arp

knitstructure+

coating

offib

ringelw

ithSM

Csandfib

roblasts

∼5m

m—

ECs,SM

Cs,and

fibroblasts

isolatedfro

movinec

arotid

arterie

s;EC

sseeding

intheinn

erlay

erusingar

otatingcham

ber;

dynamiccultu

rein

abioreactor

under

physiologicalpressurec

onditio

nsfor2

1–28

days;

histo

logicalanalys

is:ad

ense,hom

ogenou

sdistrib

utionof

cells

throug

hout

thes

caffo

ldthickn

ess,extensive

collagen

Mod

el:sheep

(6mon

ths);

ID∼6m

m;

cellu

larg

rafts

(dyn

amic

cultu

refor2

1–28

days);

at3mon

ths1

graft

with

asig

nificantsteno

sis;

otherg

rafts:com

pletely

patent,

nothrombu

sformationand

aneurysm

,infectio

n,or

calcificatio

n;remod

eling:confl

uent

endo

thelium,collagenand

otherE

CMcompo

nents,and

elastin

(not

assembled

incohesiv

esheets);

[43]


Table6:Con

tinued.

Material

Fabricationmetho

dID


Biologicalin

vitro

results

Biologicalin

vivo

results

Reference

form

ation,

andno

evidence

ofcalcificatio

n

fibrin

completely

degraded

after

1mon

th,sparseP

LDLA

materialrem

ainedat6mon

ths

PCL/PL

A:70:

30copo

lymer

of𝜀-caprolacton

eand

L-lactide;PL

DLA

:96:

4copo

lymer

ofL-lactidea

ndD-la

ctide.

SRS:sutureretentionstr

ength;BP

:burstpressure;𝐶

:com

pliance;∘

rangeo

faverage

valuesofgraft

swith

3differentm

ediacompo

sitions;H

P:hydraulic

perm

eability;UTS

:ultimatetensilestre

ngth;𝐸

:elasticm

odulus;

𝜀𝑏:strainatbreak;∗

values

wereg

raph

icallyread.

HUVEC

s:hu

man

umbilicalvein

endo

thelialcells;

HCA

ECs:hu

man

coronary

artery

endo

thelialcells;

NIH

/3T3

:murinefi

brob

lastcelllin

e;TC

Ps:tissue

cultu

replates.


Silastic tubing

Myofibroblast layer

Mesothelium

(a) (b)

Figure 3: Sketch of the in vivo bioreactor approach; (a) fibrous capsule formed around a Silastic tubing in the peritoneal cavity; (b) afterremoval of Silastic tubing, reversion of fibrous capsule, the mesothelium was the inner layer.

SMC

Fibroblasts

TEVM

TEVA

SMC

Fibroblasts

TEVTTEVTEVEV

TEV

(a)

stdTEVMA

SMC

Fibroblasts

(b)

ssTEVMASMC Fibroblasts

(c)

Figure 4: Sketch of different cell sheet-based approaches; (a) tissue-engineered vascular media (TEVM) and tissue-engineered vascularadventitia (TEVA) were assembled by rolling a single cell sheetof SMCs or fibroblasts; (b) the standard approach, to produce aTEVMA, required rolling a fibroblast sheet around a SMC sheet(stdTEVMA); (c) the single-step TEVMA was produced due to thepresence of SMCs and fibroblasts in the same sheet (ssTEVMA),from [48]. The publisher for this copyrighted material is Mary AnnLiebert, Inc. publishers.

the mechanical properties, making them more similar tonative blood vessels [26, 27] (Tables 4 and 6). Specifically,compliance and burst pressure of PCL/tropoelastin scaffoldsdecreased and increased, respectively, when compared toscaffolds made of tropoelastin alone [26, 27] (Tables 4 and6).

2.6. TEVGs without Scaffolds. Some studies were focused onfabrication of completely biological TEVGswithout the use ofscaffolds. This strategy avoids problems that may be relatedto synthetic materials, such as inflammation, stenosis, andinfection, allowing for the complete graft integration andincrease of the patency rate [8, 25]. Three main approacheshave been developed: the in vivo bioreactor, the cell sheet-based, and the cell ring-based methods.

The in vivo bioreactor approach consists in the use of thebody as a bioreactor, as a consequence of the foreign bodyreaction due to the implantation of a “nonself ” material [6].A Silastic (silicone compound) tubing was implanted in theautologous peritoneal cavity inducing a foreign body reactionwith the formation of a fibrous capsule around the mandrel[8, 25, 80]. The fibrous capsule was composed of myofibrob-lasts, collagen matrix, and a single layer of mesothelial cells(Figure 3(a)) [8, 25, 80]. After 2-3 weeks, the explanted tubeof tissue was removed from the Silasticmandrel and reversed,so the mesothelium represented the inner layer (Figure 3(b))[25].Themain limitations of thismethod are the requirementof a double surgery and the possible TEVG adhesion to theperitoneal wall during maturation [8, 80] and, hence, thedamage to the mesothelium (the external layer).

The cell sheet-based approach consists of the in vitrogrowing of cells in the presence of ascorbic acid in the culturemedium to generate a large production of ECM [8]. Aftera maturation period, cell sheets were detached from theculture flasks and rolled onto amandrel (Figure 4) [8]. TEVGswere cultured for about 8 weeks until the layers merged ina uniform structure (Figure 4) [8]. The major limitation ofthis approach is the long time required for the in vitro TEVGgrowth prior to implantation (6 to 9 months [81]).

The cell ring-based approach consists of the preparationof TEVGs from aggregated cells and cell-derived ECM, dueto the use of annular agarose wells [73]. Immortalized rataortic smooth muscle cells seeded into annular agarose wellswith different ID (2, 4, or 6mm) aggregated and formedthick tissue rings within 2 weeks of in vitro static culture(Figure 5) [73]. Tissue rings could be fused to generate tubularconstructs. Specifically, tissue rings were placed onto siliconetube mandrels and cultured in close contact for 7 days toprepare cohesive tubular constructs [73].


Side

vie

wTo

p vi

ew

Ring-shaped agarose wells

Cell suspension in agarose wells

Cells aggregate to form rings on posts

(a)

1 cm

(b)

Side

vie

w

Polycarbonate mold PDMS template Agarose well

w

p

h

(c)

Figure 5: Sketch of cell ring-based approach; (a) schematic of cell seeding and ring self-assembly process; monodisperse cells (black dots) arepipetted into ring-shaped agarose wells; within 24 hours, cells form ring-shaped aggregates (black circles) contracted around center posts; (b)photograph of 2mm diameter rat smooth muscle cell rings (arrowhead) in agarose wells; (c) schematic of agarose well fabrication process;silicone elastomer (PDMS) is cured on machined polycarbonate molds to form a template; molten agarose is poured into the PDMS templateto form agarose wells; polycarbonate mold dimensions (post diameter, 𝑝; well width,𝑤; and well height, ℎ) can be modified to make cell ringsof different sizes (courtesy of Dr. Marsha W. Rolle, Worcester Polytechnic Institute).

The most relevant and recent (2008–2013) studies onTEVG development without the use of scaffolds are reportedin Table 7, focusing on the fabrication method and mechan-ical and biological performances. Cell sheet-based TEVGs,developed by L’Heureux’s group and involved in a clinicaltrial, are widely described in Section 4.

3. Clinical Trials of VascularTissue Engineering

Clinical trials were conducted by Shin’oka’s et al. [52] andL’Heureux’s [82] groups using two different types of TEVGs.Although Shin’oka’s group did not develop TEVGs for smalldiameter blood vessel regeneration (ID = 12–24mm), theirstudies are described in this paragraph because they carriedout the first successful clinical trial of TEVGs.

In 2001, Shin’oka et al. [52] performed the first successfulclinical trial of a TEVG in humans (25 patients), specificallyin the high-flow low-pressure pulmonary venous systemof pediatric and young patients. The patients were nottreated with long-term antiplatelet and anticoagulation ther-apy [83]. The investigated TEVGs consisted of a polymericscaffold made of PGA or PLLA fiber-based meshes coatedwith a 50 : 50 copolymer of L-lactide and 𝜀-caprolactone(PCLA/PGA or PCLA/PLLA) [83]. Specifically, TEVGs werefabricated by pouring and freeze-drying a PCLA solutiononto the PGA or PLLA woven mesh (ID = 12–24mm) [14].Shin’oka’s group designed these scaffolds to improve the poorcompliance match with blood vessels and the poor surgicalhandling properties of PGA fibers [83]. For the first surgeries,grafts were in vitro seeded with autologous vascular cellsisolated from peripheral vein biopsies [52, 83, 84]. This pro-cedure was time-consuming, limiting its clinical utility: cellexpansion took approximately 60 days and cell cultured onto

scaffolds approximately 10 days, increasing the contaminationrisks. In addition, this methodology demanded an invasivetechnique for cell harvesting and showed the difficulty tocollect healthy cells fromdiseased patients [83, 84].Therefore,another cell type was found to avoid the limitations related tothe use of autologous vascular cells. The improved procedureconsisted of in vitro scaffold seeding with autologous bonemarrow-derived mononuclear cells, isolated from the iliaccrest, instead of autologous vascular cells [84]. Bone marrowmononuclear cells could be harvested the same day of thevascular surgery and required a short scaffold incubationtime before the implant (2 hours) [9, 84]. Approximately2 years after implantation, TEVGs were functional, withoutany sign of thrombosis, stenosis, obstruction, aneurysm, orcalcification [4, 9]. Approximately 7 years after implantation,the long-term followup showed no mortality related toTEVGs. Nevertheless, 16% of patients showed TEVG stenosisthat was successfully treated with angioplasty and stenting[4, 9, 14, 83, 84]. This complication was developed in TEVGswith smaller diameter (ID < 18mm) [83]. The clinical trialdemonstrated the TEVG growth potential and also someTEVG limitations: long-term graft stenosis [4, 9, 14, 83, 84]and lack ofmechanical properties necessary for a use as high-pressure arterial replacements [5].

L’Heureux’s group [82] developed a cell sheet-basedtechnology to generate completely autologous TEVGs andperformed a clinical trial for hemodialysis in 10 patients withend-stage renal disease and previous hemodialysis failure.TEVGs were implanted as arteriovenous shunts in the limbthat was contralateral to the existing graft, specifically placedin the upper arm, typically between the humeral artery andaxillary vein [81]. To fabricate completely autologous TEVGs,dermal fibroblasts were isolated from a skin biopsy and in


Table7:Stud

ieso

fthe


(2008–2013)o

nTE

VGdevelopm

entw

ithou

tthe

useo

fscaffo

lds.

Fabricationmetho

dCelltype

IDMechanicalproperties

Biologicalin

vivo

results

Reference

Invivo

bioreactor

approach:

device

implantedin

the

periton

ealcavity

toattract

cells

arou

ndatub

ular

scaffold;

scaffold=str

etchableurethane

+silicon

ewith

acoatin

gof

nonadh

esivem

aterial

Autologous

tissue

6mm

outer

diam

eter

Axialtensiletests

UTS

=0.2[0.04,0.45]bMPa,

𝜀𝑏=0.75

[0.37

,0.86]

b ;circumfer

entia

ltensiletests

UTS

=0.07

[0.02,0.72]bMPa,

𝜀𝑏=0.69

[0.39

,1.09]

b ;sim

ilartona

tiveo

vine

bloo

dvessels:

axialtensiletests

UTS

=5.7[3.39

,7.12

]bMPa,

𝜀𝑏=0.9[0.57,1.2

5]b ;

circumfer

entia

ltensiletests

UTS

=4.43

[2.76

,5.54]

bMPa,

𝜀𝑏=0.72

[0.65,1.15]

b

Mod

el:sheep;

device

implantatio

nfor10days:

manypo

rous

spaces

inneotissue;

somec

ontractilep

rotein

expressio

n;no

prod

uctio

nof

collagenbu

ndles

andela

stin;

carotid

artery

interposition:

quickrupture;

patch

esin

carotid

artery

(2we

eks):1

graft

failedaft

er1w

eek;

otherg

raftremainedpatent;

aneurysm

form

ationin

both

cases

[72]

Cellrin

g-basedapproach:

seedingof

cells

into

annu

lar

agarosew

ells+cell

aggregation+form

ationof

thicktissuer

ings

(static

cultu

re);

tissuer

ings

couldbe

fusedto

generatetubu

larc

onstr

ucts

Rataortic

SMCs

2,4,and

6mm

Circum

ferentia

ltensiletests

ID2m

mday8

:UTS

=169±45

kPa,

𝐸=0.81±0.28

MPa,𝜀𝑏∼0.46∗

;day14:UTS

=97±30

kPa,

𝐸=0.50±0.09


;ID

4mm

day8

:UTS

=339±131k

Pa,

𝐸=1.2

1±0.46


;day14:UTS

=201±

63kP

a,𝐸=0.71±0.20


;ID

6mm

day8

:UTS

=503±76

kPa,

𝐸=1.9

8±0.4M

Pa,𝜀𝑏∼0.46∗

;day14:UTS

=302±42

kPa,

𝐸=1.0

8±0.14MPa,𝜀𝑏∼0.47∗

—[73]

Cellsheet-b

ased

approach:

prod

uctio

nof

acom

pletely

autologous

TEVG

(7mon

ths)

+air-drieddevitalization+

freezing+sto

ringfor5

mon

thsa

t−80∘

C+seeding

with

autologous

ECs(cultu

refor4

days)

Autologous

fibroblastsand

ECsisolatedfro

mskin

and

vein

biop

sies

4.8m

mBP

=64

07±633m

mHg,

SRS=261±

20gf,

𝐶=3.1±

0.7%

Clinicaltrial(1p

atient);

redu

ctionof

thep

rodu

ctiontim

e(∼2weeks

versus∼6–

9mon

ths);

implantedwith

outcom

plication;

noevidence

ofleakagea

tanastomoses;

at8weeks:

nocomplication;

graft

patency;

stablediam

eter;

noadverser

eactions

[74]


Table7:Con

tinued.

Fabricationmetho

dCelltype


Biologicalin

vivo

results

Reference

Cellsheet-b

ased

approach

(singles

tepmetho

d):tun

ica

mediaandadventitia

prod

uced

byrolling

atissue

sheetcon

tainingbo

thcell

typesc

ontig

uously(SMCs

grew

besid

efibrob

lasts

inthe

samec

ulture

dish)

SMCs

isolatedfro

mhu

man

umbilicalcord;

TEVMAprod

uced

bySM

Csandeither

DFs

(TEV

MA-

DF)

orSV

Fs(TEV

MA-

SVF)

4.5m

m

TEVMA-

DF

BP∼1000∗

mmHg;

circumfer

entia

ltensiletests:

UTS∼2.1∗MPa,𝐸∼12∗

MPa,

𝜀𝑏∼30∗

%;

stress-r

elaxatio

ntests:

𝐸𝑖∼8∗

N,𝐸𝑒∼6∗

N;

TEVMA-

SVF

BP∼250∗

mmHg;

circumfer

entia

ltensiletests:

UTS∼0.8∗

MPa,𝐸∼2∗

MPa,

𝜀𝑏∼35∗

%;

stress-r

elaxatio

ntests:

𝐸𝑖∼5∗

N,𝐸𝑒∼2∗

N

—[48]

Cellsheet-b

ased

approach

(singles

tepmetho

d):tun

ica

mediaandadventitia

prod

uced

byrolling

atissue

sheetcon

tainingbo

thcell

typesc

ontig

uously(SMCs

grew

besid

efibrob

lasts

inthe

samec

ulture

dish)

SMCs

andfib

roblastsiso

lated

from

3distincthu

man

umbilicalcords;

aTEV

MA-

arteria

lTEV

MA;

vTEV

MA-

veno

usTE

VMA

4.5m

m

aTEV

MA

BPe∼

250∗

mmHg;

circumfer

entia

ltensiletests:

UTS∼3∗

MPa,𝐸∼2∗

MPa,

𝜀𝑏∼30∗

%;

stress-r

elaxatio

ntests:

𝐸𝑖∼2∗

N,𝐸𝑒∼1.5∗

N;

vTEV

MA

BPe∼

100∗

mmHg;

circumfer

entia

ltensiletests:

UTS∼1∗MPa,𝐸∼0.5∗

MPa,

𝜀𝑏∼30∗

%;

stress-r

elaxatio

ntests:

𝐸𝑖∼1∗

N,𝐸𝑒∼0.5∗

N

—[75]

Cellsheet-b

ased

approach

+decellu

lariz

ation(dTE

VM):

sheetfrom

fibroblasts(21

days)+

decellu

lariz

ationby

osmoticshock+seedingof

SMCs

+cultu

refor7

days

+cellu

larsheetrolledon

amandrel+cultu

refor2

1days

Hum

anarteria

lSMCs

isolated

from

anum

bilicalcord;

DFs

obtained

from

areductiv

ebreastsurgery;

dTEV

M-D

Fob

tained

from

DFs;

dTEV

M-SVF

obtained

from

SVFs;

TEVM

standard

cell-sheet

graft

s(no

decellu

lariz

ation),

prod

uced

bySM

Cs

4.5m

m

dTEV

M-D

FUTS∼3.1∗MPa,𝐸∼22∗

MPa,

𝜀𝑏∼23∗

%;

BPe∼

2600∗

mmHg;

dTEV

M-SVF

UTS∼2.2∗

MPa,𝐸∼20∗

MPa,

𝜀𝑏∼20∗

%;

BPe∼

1800∗

mmHg;

TEVM

UTS∼0.8∗

MPa,𝐸∼5∗

MPa,

𝜀𝑏∼3∗

%;

BPe∼

400∗

mmHg;

UTS

,𝐸,and𝜀𝑏calculated

from

circum

ferential

tensile

tests

—[76]


Table7:Con

tinued.

Fabricationmetho

dCelltype


Biologicalin

vivo

results

Reference

Microtissue

self-assembly

approach:m

icrotissues

prod

uced

high

eram

ountso

fEC

M+bioreactor

toassemble

microtissues

inatub

ular

shape+

dynamiccultu

refor14

days

(pulsatilefl

owand

circum

ferentialm

echanical

stim

ulation)

Microtissues

compo

sedof

myofib

roblastsandEC

s;HAFs

andHUVEC

susedfor

microtissueg

eneration

3mm

——

[77]

Scaffold-fre

erapid

prototyping

bioprin

tinga

pproach:cells

aggregated

into

cylin

ders+

biop

rintin

gof

cellcylin

ders

usingagaroser

odsa

smolding

template+

2–4days

inthe

mold(fu

sionperio

d)+

maturationin

aperfusio

nbioreactor

HUVS

MCandHSF

aggregated

into

cylin

ders

0.9to

2.5m

mou

ter

diam

eter

——

[78]

TEVMA:tiss

ue-eng

ineeredvascular

mediaandadventitia;TE

VM:tiss

ue-eng

ineeredvascular

media.

UTS

:ultimatetensiles

treng

th;𝜀𝑏:strainatbreak;𝐸:elasticm

odulus;𝐸𝑖:initia

lmod

ulus;𝐸𝑒:equ

ilibrium

mod

ulus;B

P:bu

rstp

ressure;BP

e:estim

ated

burstp

ressure;∗

values

wereg

raph

icallyread;bdataexpressed

asmedians

(min.,max.),

SRS:suture

retentionstr

ength,C:

compliance.

SVFs:h

uman

saph

enou

sveinfib

roblasts;

DFs:h

uman

derm

alfib

roblasts;

HAFs:p

rimaryhu

man

artery-derived

fibroblasts;

HUVEC

s:hu

man

umbilicalvein

endo

thelialcells;

HUVS

MCs

:hum

anum

bilicalvein

smoo

thmuscle

cells;H

SFs:hu

man

skin

fibroblasts.


vitro cultured with ascorbic acid, facilitating the productionof extracellular matrix proteins [5, 81]. Approximately 6weeks after the seeding, fibroblast sheets were detached fromculture flasks, wound around a temporarymandrel (diameter= 4.8mm), and led to maturation for approximately 10 weeksto generate a homogenous tissue [5, 81]. Then, autologousendothelial cells isolated from peripheral vein or endothelialprecursors taken from circulating blood were seeded intothe TEVG inner layer devitalized by air-drying, few daysbefore surgery [5]. TEVGs required 6 to 9 months of invitro growth prior to implantation [81]. Primary patency wasmaintained in 7 of the remaining 9 patients 1 month afterimplantation and in 5 of the remaining 8 patients 3–6monthsafter implantation [81]. There were 3 TEVG failures due tothrombosis or aneurysm during the safety phase that wasused to evaluate TEVGmechanical stability. In spite of the useof completely autologous grafts, one patient showed a TEVGaneurysm, mediated by an acute immune response that wascaused by high levels of immunoglobulin G present in aserum batch during the cell culture [8, 81]. At implantation,TEVGs exhibited a poor compliance (3.4 ± 1.6%/100mmHg,experimental evaluation) in comparisonwith human internalmammary artery; however, 6 months after implantationthe TEVG compliance increased and was equal to 8.8 ±4.2%/100mmHg (measured by Doppler ultrasound) [85].Some challenges are related to L’Heureux’s clinical trial;participants were end-stage renal disease patients andTEVGsare subject to repeated puncture and great hemodynamicloads typical of the hemodialysis access [5, 8]. To date,L’Heureux’s group intends to decrease the cost (≥$15,000 pergraft [23]), the complexity, and the time for the productionof their TEVGs [8]. They performed a first human implantof a devitalized frozen graft in a hemodialysis patient: 5 daysprior to surgery, the TEVG was rehydrated and its lumenwas seeded with autologous endothelial cells [8]. Fivemonthsafter implantation the TEVG worked with no evidence offailure [8].

4. Conclusions and Future Perspectives

This review is focused on the most relevant and recentadvances concerning vascular tissue engineering for smalldiameter blood vessel applications, specifically the develop-ment of TEVGs made of decellularized matrices and naturaland/or biodegradable synthetic polymers and their realiza-tion without the use of scaffolds. Particular attention is paidto scaffolds made of natural and/or biodegradable syntheticmaterials, highlighting their advantages and disadvantagesand comparing mechanical and biological performances ofnatural and synthetic materials.

Main failures of ePTFE prosthesis were related to neoin-timal hyperplasia, remarkable calcification, thrombosis, andforeign body reaction [60]. During clinical trials or in vivoanimal tests, TEVGs often showed the same failure causes ofePTFE prostheses (Tables 3–7 and Section 4). The challengeof vascular tissue engineering still remains the developmentof TEVGs that are able to regenerate a functional endothe-lium and exhibit mechanical properties similar to nativeblood vessels. Although completely autologous TEVGs have

reached promising results in clinical trials, the developmentof readily available TEVGs characterized by an economicalfeasibility still remains a major challenge. In fact, the fab-rication of completely biological TEVGs without the use ofscaffolds requires long time and high cost of manufacturing(Section 2.6).

Among the different approaches based on the use ofscaffolds, the direct implantation of cell-seeded scaffoldswithout the use of bioreactors simplifies the procedure andreduces the cost related to the use of bioreactors.Thismethodis based on the concept that the cells seeded in scaffold play akey role in the recruitment of the host cells for in vivo tissueregeneration, as shown by Shin’oka’s group. The implanta-tion of acellular scaffolds for in situ tissue regeneration bythe host would avoid the limits correlated to the in vitrocellularization of scaffolds, specifically the cell choice andexpansion, and the use of a bioreactor that is time-consumingand technical complex due to the mandatory application ofGood Manufacturing Practice (GMP, EudraLex—Volume 4)for the clinical use. Therefore, both of these approaches leadorwould lead to off-of-the-shelf scaffolds for the regenerationof small diameter blood vessels, considerably simplifying theprocedure for fabrication of TEVGs and reducing the cost andthe time of manufacturing.

Among all the available materials for small diameterblood vessel regeneration, hybrid scaffolds exhibit promisingand tailorable results due to the combination of appro-priate mechanical properties of synthetic polymers withexcellent biological performances of natural polymers (Sec-tion 2.5). Furthermore, an important perspective wouldbe the direct introduction of antiinflammatory and anti-proliferative drugs in TEVGs, due to the better knowledge ofmolecular pathogenesis of artery disease.

Conflict of Interests

The authors declare that there is no conflict of interestsregarding the publication of this paper.

Acknowledgments

The work was supported by Cariplo Foundation funds 2007to G. Freddi (Project no. 2007-5457). The authors would liketo thank Dr. Marsha W. Rolle for the useful discussion andcomments.

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