The Roles of Tissue Engineering and Vascularisation in the Development of Micro-Vascular Networks- A Review

pital Hampstead NHS Trust, London NW3 2QG, UK

Received 1 April 2004; accepted 5 July 2004

Available online 20 August 2004

Keywords: Micro-vascular; Tissue engineering; Capillary beds; Tissue vascularisation

3.1.2. Conguration of vessels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1863

ARTICLE IN PRESS3.2. Extracellular matrix scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1863

3.2.1. Biomaterial substitutes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1864

3.2.2. Superstructure . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1867

Corresponding author. Tel.: +44-20-7830-2901; fax: +44-20-7472-6444.E-mail address: [email protected] (A.M. Seifalian).0142-9612/$ - se

doi:10.1016/j.biContents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1858

1.1. Reconstructive surgery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1858

1.2. Tissue engineering and its limited success . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1858

1.3. Microvessel development. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1858

1.4. Tissue perfusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1859

2. Search methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1860

3. Tissue-engineered microvessel beds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1860

3.1. Micro-vascular tissue engineering. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1860

3.1.1. Types of prostheses . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1861Abstract

The construction of tissue-engineered devices for medical applications is now possible in vitro using cell culture and bioreactors.

Although methods of incorporating them back into the host are available, current constructs depend purely on diffusion which

limits their potential. The absence of a vascular network capable of distributing oxygen and other nutrients within the tissue-

engineered device is a major limiting factor in creating vascularised articial tissues. Though bio-hybrid prostheses such as vascular

bypass grafts and skin substitutes have already been developed and are being used clinically, the absence of a capillary bed linking

the two systems remains the missing link. In this review, the different approaches currently being or that have been applied to

vascularise tissues are identied and discussed.

r 2004 Elsevier Ltd. All rights reserved.bDepartment of Plastic Surgery, Royal Free HosBiomaterials 26 (2005) 18571875

Review

The roles of tissue engineering and vascularisation in the developmentof micro-vascular networks: a review

Ruben Y. Kannana, Henryk J. Salacinskia, Kevin Salesa, Peter Butlerb,Alexander M. Seifaliana,

aBiomaterials & Tissue Engineering Centre (BTEC), University Department of Surgery, Royal Free and University College Medical School,

University College London, Rowland Hill Street, London NW3 2PF, UK

www.elsevier.com/locate/biomaterialse front matter r 2004 Elsevier Ltd. All rights reserved.

omaterials.2004.07.006

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4. Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1870

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more case-specic [4]. More recently, the introduction of was seeded with cultured keratinocytes. The CSR was

INateriperforator aps has improved the utility of aps as areconstructive option [5]. Nevertheless, free tissuetransfer confers morbidity to the patient as one functionis sacriced for another [6]. Therefore, alternativesources of tissue for reconstruction are needed.

1.2. Tissue engineering and its limited success

More than 21 million patients per year in the USalone rely on biomedical implants [7]. These include

used in 12 pediatric burn wounds and by 14 days45.7714.2% had vascularised as compared to 9871%in split skin grafts [15]. However, no clinical studies onscaffolds with an inherent vascular network or incorpo-rated angiogenic factors have been done.

1.3. Microvessel development

Human microvasculature begins with arteries dividingconsecutively into smaller branches like meta-arterioles4.1. Future perspectives. . . . . . . . . . . . . . . . . . . . . . .

References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

Further-reading . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1. Introduction

1.1. Reconstructive surgery

In current practice, reconstructive surgery is essentialin order to achieve wound closure following majorcancer resections and trauma. Traditionally, reconstruc-tive surgeons have employed the reconstructive ladderconcept; a hierarchical system by which the simplestoption, usually skin grafting or primary closure, ischosen rst. If this fails, then a slightly more complexprocedure is selected. With the increasing demand forquality wound closure, maximal aesthetics with mini-mum morbidity, the concept of the reconstructiveelevator [1] is applicable. As opposed to the reconstruc-tive ladder, the best option, which in most cases is a ap,is chosen initially.A ap may be dened as a segment of tissue with an

independent blood supply. These may be classied intolocal, regional and free aps. Free aps are based on theconcept of angiosome [2]; an area of tissue with aninherent vascular network supplied by a single vascularpedicle. Buncke and McLean performed the rst micro-vascular tissue transplant in 1969 [3]. Since then, freeaps have revolutionised reconstructive surgery as theyare more versatile and reliable than other aps. Flapsare now being pre-fabricated prior to surgery to become3.2.3. Cells seeded onto matrices . . . . . . . . . . .

3.3. Vascularisation . . . . . . . . . . . . . . . . . . . . . . . . .

3.3.1. Inosculation. . . . . . . . . . . . . . . . . . . . . .

3.3.2. Angioinduction . . . . . . . . . . . . . . . . . . .

ARTICLER.Y. Kannan et al. / Biom1858bypass grafts, dental implants and articial dermalsubstitutes like Integras. Current technology limits thesurvival of these implants as they depend initially upondiffusion and in the later stages on neovascularisation.Overdependence on the former limits the thickness of. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1870

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1870

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1875

these implants, while depending on neovascularisationcan cause excessive brovascular ingrowth and hencescarring. There are no tissue-engineered constructspresently available which have an inherent vasculaturesuch as a capillary bed ready to be connected to the hostvascular system. This represents a major limitation asthe driving force of living tissue is vascularity. Althoughattempts to develop small-diameter vascular grafts havebeen successful in vitro [8,9], it remains to be seenthrough clinical studies whether they are reliable enoughto sustain ow in vivo [1012]. Even so, they are limitedby the absence of a viable capillary network for nutrientexchange. This missing link represents the bridgebetween the host and the tissue-engineered implant.Most clinical studies in articial tissue so far have

utilised dermal substitutes like Integras in the treatmentof burns. After application, it undergoes imbibition,broblast migration, neovascularisation and nallyvessel maturation. This process takes months, afterwhich skin grafts or cultured keratinocytes may beapplied. One centre used it in 30 patients with goodresults [13], while others have successfully seededcultured Fb and keratinocytes onto it prior to clinicalapplication. They found superior results when it wasdone with concurrent skin grafting [14]. Sheridan andco-workers developed a composite skin replacement(CSR) composed of acellular allogeneic dermis, which. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1867

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1868

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1868

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1869

PRESSals 26 (2005) 18571875(80100 mm) until nally forming capillaries (1015 mm).These vessels serve to redistribute blood and its nutrientswhilst lowering the pressure head. This allows blood toperfuse the tissue, allowing more efcient exchange ofmetabolites. Eventually, the capillaries unite into

post-capillary venules (PCVs), venules and nally veinsfor the return of waste products. The venular componentof vasculature also serves as capacitance vessels [16].The vascular tree is formed during the early gestation.

Angiogenic cells form clusters which coalesce to formsolid tubes, which eventually canalise to form bloodvessels [17]. The outer ring consists of angioblasts whichform the vessel walls. The subsequent differentiation ofthese precursor angioblasts into endothelial cells (EC)and the denovo formation of a vascular network aretermed vasculogenesis. These vessels are capillary-liketo begin with and eventually differentiate into eitherarteries or veins [18]. The creation of this frameworkdepends on guidance molecules within the matrix suchas ephrins at the arterio-venous interface [19] andreversion-inducing cysteine-rich (RECK) protein forparacellular proteolysis [20].The adult vascular network remodels itself by

arteriogenesis with the opening up and then thesubsequent enlargement of existing collaterals (so calledcollateral enlargement), as well as the formation of

Remodelling is mediated by monocytes and endothe-lial progenitor cells (EPCs) as shown in Fig. 1 and is a

ability of the membrane (P) and surface area (S) as

ARTICLE IN PRESSR.Y. Kannan et al. / Biomaterials 26 (2005) 18571875 1859completely new vessels from the already existing vessels(so-called arterialisation) [21]. Micro-vascular remodel-ling is a mechanistic process delineated to specic tissuetype and specic stimuli. Therefore, with the exceptionsof skeletal muscle responding to exercise and the femalemenstrual cycle itself, micro-vascular remodelling islimited solely to pathological situations, in particularinammation, wound healing, ischemia, and that ofhypoxia, together with a few other rare circumstances.Fig. 1. Series of events illustrating the dual nature of vessel

development: angiogenesis and arteriogenesis. Keys: EPC, endothelial

progenitor cells; BM, basement membrane; SM, smooth muscle cells.represented below:

D PSC:P in turn, depends on effective pore area (A) and porelength (L). Capillary exchange is a mixture of diffusion-and ow-limited exchange, with the latter beingimportant for the transfer of water-soluble moleculesand the former for gas transport.Plasma is a two-phase uid consisting of solvent and

solute phases. While solvents pass unimpeded into theextracellular matrix (ECM), solutes are transported bymeans of convection (bulk transport) and diffusion [28],represented by the formula

F QD;mechanism that is dependent on changes in shear stress.Studies have shown that turbulent ow at low ow ratesof 1.5 dynes/cm2 itself activates ECs while a normallaminar shear stress of 815 dynes/cm2 does not [22]. Ina rabbit femoral artery occlusion model, collateralrecruitment occurred within a week, with removal ofunwanted vessels by 3 weeks [23]. Angiogenesis is thecoordinated migration and proliferation of EC andpericytes from the existing vascular bed [24] and theirsubsequent maturation and stabilisation by envelopingsmooth muscle cells (SMCs) [25]. Stimulated byhypoxia, these vessels proliferate by either capillarysprouting [21] or intussusception [26], particularly intovenules. Vasculogenesis occurs from migrating dediffer-entiated EPCs which form tubules [27].

1.4. Tissue perfusion

Microvessels divide into numerous smaller branchesover a given volume of tissue, thus maximising theavailable area for nutrient exchange. In this microenvir-onment, stasis of ow is prevented by the Fahraeus-Lindqvuist effect, repulsive charges between blood cellsand vessel wall, as well as the thin glycocalyx lm on theendothelial layer (Fig. 2) [16].As blood slows down in the meta-arterioles, intra-

vascular transit time increases. As this happens, nutrientexchange shifts from ow-limited to diffusion-limitedexchange within the capillaries (Fig. 3). Diffusion (D) isdependent on the concentration gradient (C), perme-Fig. 2. Factors preventing thrombogenicity in low-ow states.

3. Tissue-engineered microvessel beds

3.1. Micro-vascular tissue engineering

Microvessel beds provide the vascular infrastructurefor living tissue [6]. The development of independentlyvascularised articial tissue requires a source of vascu-larisation (intrinsic existing blood vessels) and a matrixthrough which exuding nutrients can perfuse cells [30].The matrix functions as a reservoir of growth factors toinduce incoming blood vessel (angioinduction) as well asa scaffold to seed cells like EPCs, which participate inthe formation of newer blood vessels by the process ofarteriogenesis (Fig. 4). Microvessel networks, articialmatrices and neovascularisation constitute the triad oftissue vascularisation (Fig. 5).A successful micro-vascular system depends on the

graft material and how it is arranged. The materialwould have to be non-thrombogenic and must havesimilar compliance to native vessels to avoid intimal

IN PRESSaterials 26 (2005) 18571875where F is solute ow and Q is convective ow. Soluteefux is dependent on the rate of solvent transfer withinthe medium. Solutes like glucose are transported viaconvection with water (solvent), whilst gaseous ex-change is based on Kroghs diffusion model.This is augmented by the 5mmHg negative pressure

as a result of ECM hydraulic conductivity, the ability oftissue to conduct water across it. Hydraulic conductivity(H) of ECM is proportional to the hydraulic radius ofthe gel (the ratio of its porosity, P, to the net surfacearea of the glycosaminoglycans/GAGs):

Har=S:

Fig. 3. Flow- and diffusion-limited exchange.ARTICLER.Y. Kannan et al. / Biom1860The interstitial conductivity of ECM is proportional toits hydration and is usually in the range of10101013 cm3/s/dyne. This force is balanced by theintrinsic hydraulic resistivity of the ECs lining themicrovessels. Studies have shown this to be in thevicinity of 108 cm3/s/dyne [29]. Nutrient exchangedepends on the balance between hydrostatic, osmoticand interstitial pressures [16].

2. Search methods

All the studies were identied by following databases:Pubmed; Ovid online; Infotrieve; Proquest; Sciencedirect; Isi web of science; Biomed central; Ingenitaselect; Elsevier texts and Blackwell-synergy searchesbetween years 1966 and 2003, with the followingkeywords: Tissue-engineered blood vessels; Vascularisa-tion of acellular tissue; Vascularisation of tissue;Angiogenesis; Vascular tissue engineering; Microvascu-lature; Tissue-engineered aps; Pre-fabricated aps; Softtissue engineering; Polymers in vascular tissue engineer-ing; Capillary beds; Microvasculature; Polymers intissue engineering; Collage; Dacron; PTFE; Alginate.hyperplasia [31] at its arterial end and must besufciently porous to allow nutrient exchange at thecapillary level. An articial capillary network wouldinclude small arteries (12mm) conducting blood intoan arteriolar network (1001000 mm), which wouldeventually end in capillary-like vessels of 1015 mm.These end capillaries would need to be within150200 mm of every target cell [30] before converginginto a venous collecting system. Therefore, botharteriolar and venular components of the scaffold wouldhave to be in close proximity.Fig. 4. Inter-relationships between various methods of vessel devel-

opment. Keys: EC, endothelial cell; EPC, endothelial progenitor cells.

3.1.1. Types of prostheses

The following section has been subdivided intomacro- and micro-vascular grafts, the reason being thatmicro-vessels are being developed from previous tech-

nology as used in the development of larger macro-vessels, and only rarely are completely new material-based approaches used to date. This discussion isnecessary to deal with the issues of potential overlapsand discrepancies during the transition between thehigh-ow arterial system and the lower-ow capillarysystem and how, in light of these, tissue engineering ofmicro-vessels differs. The various graft options aresummarised in Table 1.

3.1.1.1. Macrovascular grafts. The current resultswith PTFE and Dacront are clinically acceptable inperipheral bypass grafts [32], where as patency issignicantly lower in grafts 6mm or smaller [3335].The primary problem at lower ow rates is thrombo-genicity, especially with Dacron-based grafts, and theincreased susceptibility of these vessels to intimalhyperplasia. Luminal modication of these grafts usingirradiation, presealing and antiplatelet/anticoagulant/growth factor incorporation [30,47,48] to reducethrombogenicity had been employed [46], but endothe-lialisation is signicantly superior [49,50]. Endothelia-lisation of vascular prostheses is dependent on thebiomaterial used and pore morphology. Matsuda andcolleagues reported that pores between 18 and 50 mm in

ARTICLE IN PRESS

Table 1

Current status of macro- and microvascular grafts

Authors Graft type Subjects ID (mm)

Harris et al. (2002) [37] PTFE Rats o1

Demiri et al. (1999) [36] PTFE Rats 1

Seifalian et al. (2003) [41] CPU Dogs 5

Shinoka et al. (1998) [176] Polyglactin-

PGA

Sheep 15

Shum-Tim et al. (1999) [177] PHA-PGA Sheep 7

Meinhart et al. (2001) [32] EC-seeded

ePTFE

Humans 6 to 7

Lambert et al. (1999) [178] Heparin-

coated

Dacron

Humans 6 to 7

Devine et al. (2001) [179] Heparin-

bonded

Dacron

Humans 7 to 9

Fig. 5. The triad of tissue vascularisation.

R.Y. Kannan et al. / Biomaterials 26 (2005) 18571875 1861Weinberg et al. (1986) [180] Collagen-Fb-

SMC

In vitro Unknown

LHeureux et al. (1998) [65] Collagen-EC-

SMC-Fb

In vitro 4.6

Wilson et al. (1995) [63] AAM Dogs 3 to 4Teebken et al. (2001) [35] AAM Pigs 5

Keys: AAM, allogeneic acellularised matrix; ePTFE, expanded polytetra

polyurethane graft; ID, internal diameter; PHA, polyhydroxyalkanoate; PGVessels Patency rates Comments

Supercial

epigastric

20% at 4

weeks

Vein grafts had 100% patency

Femoral 25% at 4

weeks

Vein grafts had 100% patency

Aorto-iliac 100% at 3.5

years

Human trials under way

Pulmonary 100% at 7

weeks

Good tissue ingrowth but increase

in vessel ID

Aorta 100% at 21

weeks

Outer PHA ring prevented

increase in vessel ID

Infra-popliteal 74% at 7

years

Comparable to vein grafts

Infra-popliteal 58% at 2.5

years

Better results than plain Dacron

Femoro-

popliteal

55% at 3.5

years

This trial shows better results than

PTFE

92% EC coverage but weak

bursting strength

Comparable bursting strengths to

human vessels

Coronary 44% at 6 All failures due to acute graftmonths rejection

Carotid 38% at 4

months

Disappointing results due to graft

rejection

uoroethylene; Dacron, polyethylene tetraphthlate; CPU, compliant

A, polyglycolic acid.

An in vivo study in a rat femoral artery model showeda 2025% patency rate at 1 month with PTFEmicrovessels (o1mm), while all vein grafts in similarsettings remained patent [36,37]. In order to improvepatency rates, luminal endothelialisation has beenattempted. Kidd and colleagues, through 1mm inter-position ePTFE vascular grafts in rat aortas, coatedwith ECM (such as matragel) and incorporated withnon-tumorigenic human keratinocyte cell line (HaCaT)or EpsteinBarr Virus (EBV)-transformed human B-celllymphocyte (TOU II-4) cell lines, showed ablumenalvascularisation with the formation of an EC lining 5weeks post-implantation [10]. Correlation of theseresults to large animal models is hindered by the factthat rat animal models rapidly endothelialise, whilst thelatter do not to the same extent, this warranting furtherresearch in this particular eld [52,9].An in vitro biological micro-vascular model has been

developed by Neumann and co-workers, wherein 80 mmdiameter nylon strands were used as templates onto

IN PRESSaterials 26 (2005) 18571875diameter are optimal for endothelialisation [42].Smaller pores would elicit an inammatory reaction[43], while increasing pore sizes and porosities improvethe radial compliance of these grafts [44] and reduce thelikelihood of intimal hyperplasia.More compliant biomaterials such as poly(carbo-

nate-urea)urethane (CPU) are now being introduced[53]. CPU is currently undergoing Phase I clinical trialsas a lower limb graft and has been found to retainsimilar compliance to native vessels after implantationand exhibited little intimal hyperplasia [39,40]. There-fore, it has the potential to be used to construct thearterial conducting component of articial capillarybeds. Synthetic protein polymers cross-linked by g-irradiation represent a new generation of biopolymers.Preliminary reports suggest that they have similarelasticity as arteries, with a controllable rate ofdegradation [45]. This recent trend illustrates thelimitations of PTFE and Dacront at lower ow ratesand stresses the need for alternative biomaterials.Materials such as PTFE and Dacron have littlepotential for micro-vessel grafts due to their poorhaemodynamics caused by lack of arterial compliance.This results in thrombogenicity to the extent that inlow-ow applications (as is the care for microvessels)below 4mm their usage is predelicted by their proprietyfor blockage. Polyurethanes such as CPU due to theirsuperior haemodynamics (a result of their arterial likevisco-elasticity and honeycomb structure) have highapplicability for microvessel development and haveshown little blockage in 2mm applications (unpub-lished data).

3.1.1.2. Micro-vascular grafts. Micro-vascular graftsare those with internal diameters of 1mm or less, andmay be classied into conducting arterial and distribut-ing capillary vessels. The construction of arterialconduits is based on small-calibre vascular grafttechnology. These new grafts may be constructed byusing (1) newer polymers such as CPU, (2) by coatinggraft lumen [5457] with anti-platelet agents or cells(most commonly endothelial cells, but also possiblybroblasts or smooth muscle cells) and (3) constructingbiological or bio-hybrid grafts in vitro prior to re-implantation [38, 5961]. As these vessels divide further,they need to branch into a distributing capillary networkwith larger pore sizes and thinner vessel walls forenhanced nutrient exchange. Although autologous veingrafts, the current gold standard for micro-vascularrepairs, are compliant and non-thrombogenic, they arelimited by the need for additional vein-harvesting

ARTICLER.Y. Kannan et al. / Biom1862procedures. Furthermore, the construction of articialvascularised tissue requires an inherent vascular net-work [58]. Vein grafts are not suitable for this purpose asit is technically impossible to dissect out a capillary bedin its entirety.which SMCs in a culture medium could attach to. Aftera week, this strand was physically removed and thegrafts were perfused in a pulsatile ow chambersimulating the arterial system (Fig. 6). After 28 days inculture, these vessels were shown to have 3040 layers ofSMCs evenly distributed throughout the lumen andstrong enough to withstand its own weight. Non-absorbable nylon was preferred to absorbable cellulosesince constructs based on cellulose had irregular, unevenluminal surfaces [43]. This model has the potential toform arterioles and capillaries in vitro, which could thenFig. 6. A viable microvessel perfusion system. Modied from Ref. [66].

Key: SM, smooth muscle cell.

regeneration with predominant sprouting from the venousend. They concluded that while an AV loop quantitativelyincreased capillary and new tissue formation, ligating theloop ensured better tissue organisation [72].

3.2. Extracellular matrix scaffolds

ECM is composed of collagen, bronectin, laminin,GAGs and water [73]. Fibroblasts synthesise non-cellular components of the matrix, as shown in Table2. Using acellular dermal grafts in SpragueDawley ratgroins, matrices were pre-fabricated by vascularisingthem with the supercial epigastric vessels. Histological

IN

Fig. 7. The effect of vessel conguration on tissue vascularisation

(modied from Ref. [72]. Group Icontrol (acellular dermis); Group

IIarterio-venous loop embedded in the matrix Group IIIarterio-

venous bundle ligated within the matrix; Group IVow-through

graft).

Table 2

Non-cellular components of ECM and their function [73]

Component Function

Collagen type I Structural framework

Collagen type IV Basement membrane for EC attachment

Collagen type VII Keratinocyte attachment

Fibronectin Cell adhesion

Laminin Protein adhesion

GAGs Cytokine and growth factor binding

ateribe transplanted into living tissue. Further results usingECs in a similar model are being undertaken [66].Bio-hybrid microvessels have also been developed in

vitro using co-cultures of EC and SMC. These cells werepassed through polypropylene capillaries at 3 dynes/cm2

laminar shear stress. The cells lined the lumen of thesetubes while SMCs managed to migrate outside these150mm diameter microvessels through pores (0.5mm) toform a two separate monolayers as in normal vessels [64].

3.1.2. Configuration of vessels

The limiting factor in metabolite exchange is theextra-vascular distance needed to be traversed. Thephysiological response to this is to pair up both thearteriolar and venular components together. Betweenthem, numerous small capillaries run across tissue at across-sectional density of 1300 permm2, with an inter-capillary distance of 34 mm to form a rich perfusingnetwork. This is well within the maximum diffusingdistance of oxygen, glucose, carbon dioxide and otherwaste metabolites [16]. In studies on rat mesentery, itwas found that arterioles and venules travelling togetherformed a counter-current system responsible for auto-regulation and improved nutrient exchange [67]. There-fore, arteries and veins should be paired for greaterefcacy.A study of the temporal relationship of pre-fabricated

aps was performed by inserting a ap of tissue underrat abdominal skin. The femoral arterio-venous pediclesinduced ap neovascularisation as early as 2 weeks post-surgery, but it took up to 8 weeks for the vascularpedicle to sustain the ap alone [68]. However, in asimilar study, the ap was found to survive after itspedicle was divided after 34 weeks [69]. Studies onhybrid aps have also shown that the pedicle can bedivided by 4 weeks, provided a highly porous biomater-ial like polyethylene is used along with living tissue as aap [70,71].Tanaka and associates used the rat femoral vessel

system to determine whether loop, ligated bundles orow-through pedicles were most effective in initiatingand sustaining tissue perfusion and neovascularisation.Using a collagen I dermal implant in the rat groin, theyformed four groups. In the rst group, the implant wasplaced on a recipient bed with no vascular pedicle. In thesecond group, a vascular loop formed of either anarterial or venous graft was employed. In the thirdgroup, an arteriovenous (AV) bundle ligated at one endwas embedded within the matrix, and in the fourthgroup an AV ow-through graft was placed across thematrix (Fig. 7). Four weeks later, the matrices were

ARTICLER.Y. Kannan et al. / Biomsectioned and stained. Matrices from group II showedthe greatest increase in volume with ow rates of up to20 times normal. However, compared to the third group,it had less tissue organisation. AV ligation (group III)showed the densest vascularity and maximal tissuePRESSals 26 (2005) 18571875 1863analysis showed inltration by mononuclear cells withinthe rst 72 h with neo-vascularisation occurring betweendays 3 and 14. Simultaneously, SMC and EC were laidalong the lines of stress. Devascularised matrices(control group) had a 90% hernia rate, while those with

an intact vascular pedicle within had no evidence of ahernia (po0.01). Therefore, it was concluded thatvascularised scaffolds were structurally and functionallysuperior to non-vascularised ones [74].

3.2.1. Biomaterial substitutes

The ideal matrix should be biocompatible, porous,supportive, readily available, permeable to incomingvessels or chemicals, non-immunogenic [62], biodegrad-able, inert, have a slow-release mechanism and, lastly, beeasily administered as an injectable. Tissue-engineeredconstructs act as temporary scaffolds, which maintaincells in situ and induce neovascularisation. Eventually,the existing matrix will be replaced by the products ofnewer cells such as collagen, laminin, elastin, GAGs orbronectin. These undergo continuous remodelling [75]as depicted in Fig. 8, and ultimately reorganise intonear-normal tissue [76]. Current scaffolds can beclassied into natural and synthetic types, each with itsown advantages and disadvantages. Natural biomater-ials are characteristically biodegradable, non-immuno-genic and bio-compatible. A list of natural ECMavailable is summarised in Table 3.

3.2.1.1. Natural. Alginate is an anionic polysaccharidecomposed of b-D-mannuronic acid and a-L-glucuronicacid cross-linked in a gel by non-toxic levels of calciumions. It has been incorporated with various growthfactors like basic broblast growth factor (bFGF),vascular endothelial growth factors (VEGF) and hepar-in within microspheres [77]. Alginate gel is a hydrogelsystem that is independent of temperature changes [78]but tends to release its chemicals within 4 days [79]. Thiscan be overcome by incorporating microspheres andpeptides [80] or cross-linking with heparin [81]. Alginategels modied with arginine-glycine-aspartic acid (RGD)peptide sequences [82] and syngeneic broblasts formedstrong gels and were more tissue-like [83]. An agarosegel cross-linked with heparin has been used to deliverbasic broblast growth factor (FGF) beads to the

ARTICLE IN PRESS

ent

Agarose Agarose Solid gel

R.Y. Kannan et al. / Biomaterials 26 (2005) 185718751864Hyaluronic acid

(HA)

Glucuronic acid+N-

acetyl-D-glucosamine

(linear GAG)

Biodegradable gel,

biocompatible,

7photopolymerisable

Fibrin glue Fibrin monomers Injectable gelaFig. 8. The effect of cell traction on tissue patterns.

Table 3

Natural materials used in ECM constructs

Biomaterials Composition Characteristics

Alginate b-D-mannuronicacid+a-L-glucuronicacid

Temperature-independ

gelCollagen/gelatine Type I collagen

(atelocollagen)

Porous and permeable

Matrigel Collagen type

IV+laminin+entactin+proteoglycansInert, permeable, injectable gelStructuporcine myocardium, but delivery characteristics arestill sub-optimal [84].Hyaluronic acid (HA), a linear glycosaminoglycan

(glucuronic acid N-acetyl D-glucosamine), has goodbiocompatibility and is degraded by hyaluronidase.However, it has poor adherence and inhibits ECproliferation, while its degradation products are pro-angiogenic [85]. To improve this, chitosan, a glycosa-minoglycan (N-acetyl D-glucosamine), has been addedto HA to form a composite material in the synthesis ofendotheliased human skin equivalent [86]. Chitosanattracts neutrophils and activates macrophages [87].Microspheres lled with FGF within similar constructswere used to improve neovascularisation in rat tissue[88]. The addition of methacrylate groups to HA,followed by exposure to light of 480520 nm wave-lengths, could photopolymerise composite HA gels intoany given shape with minimal gel swelling. RGDsequences could also be added to the composite to

Comments Modications References

Burst release Microsphere/bead

incorporation, heparin

cross-linkage (controlled

release), RGD-peptide

sequence incorporation

[78,8082]

Sub-optimal release Heparin addition (increase

factor load)

[84]

Angio-inhibitory,

degradation products are

angiogenic, non-adhesive

to cells

Chitosan cross-linkage

(improves angiogenesis),

RGD sequences (improve

cell adherence),

methacrylate addition

[85]

Structurally weak Heparin cross-linkage [90](controlled release)

Low factor-loading

capacity

Heparin cross-linkage [91]

rally weakNone[94]aIdeal for local delivery.

INateriimprove cell adherence. In vitro studies on thiscomposite showed improved broblast viability andproliferation compared to standard HA gels [89].Fibrin glue derived from blood clots is useful as an

injectable medium for the local delivery of growthfactors like VEGF. Burst release is a problem that isovercome by the incorporation of heparin for bettercontrol [90]. Collagen type I is highly porous (pore sizesof 50150 mm) and has been seeded with Fb or pre-adipocytes [91]. Chondrocytes embedded within three-dimensional (3-D) collagen gels were found to be viableand capable of synthesising proteoglycans (PGs) [92].Unfortunately, loading factors onto it is difcult, butthis can be overcome by cross-linking collagen gels withheparin. This has been shown to increase the incorpora-tion of angiogenic factors sufciently to culture humanumbilical vein cells (HUVEC) in vitro [93]. ECM gel(Matrigelt) that is composed of basement membraneproteins has been used in vitro as scaffolds [94], but hasso far few clinical uses [95]. Other natural substitutesused are porcine intestinal sub-mucosa [96], acellulardermis, cadaveric fascia and amniotic membrane [97].The most signicant drawback of these ECM substitutesis that they are structurally weak with unmodulatablehydraulic conductivity.

3.2.1.2. Synthetic. Numerous polymers have beenemployed in tissue engineering, including poly(propy-lene fumarates)/PPF, poly anhydrides (PA), polyglycolicacid (PGA), polylactic acid (PLA), polycarbonates (PC),polylactone (PL), polyurethanes (PU), poly(orthoesters)(POE), polyphosphazenes (PPZ) and poly(ethylene-terephthalate) (PET) [98] and the commonest polymerin use, polyester (PE). These are being used withincreasing frequency in biomedical devices.PGA is an inelastic polyester with high crystallinity

(4650%). It is degraded by the diffusion of waterfollowed by hydrolysis of the crystalline elements of thepolymer into glycolic acid [99]. PLA is less crystallinethan PGA but more hydrophobic, making it lesssusceptible to hydrolysis. Poly(lactide-co-glycolide)(PLGA) co-polymer in the amorphous form is a high-porosity, open-pore system which degrades in 26months by means of hydrolysis of the ester bonds[100]. A signicant drawback is that angiogenic factorsincorporated within it tend to leach out. This can berectied using microsphere, beads or bioactive foams[101] for slower release of chemicals. Experiments usingmicrospheres containing recombinant VEGF withinPLGA foams, which were then implanted into ratfascial aps, showed a more controlled release of growth

ARTICLER.Y. Kannan et al. / Biomfactors [102].1,3-trimethylene carbonate (TMC) and D,L-lactide

(DLLA) (19:81, molecular weight ratio) are amorphousco-polymers of PLA. Salt-leaching these polymersproduced foams of high pore interconnectivity andporosity with an average pore size of 100 mm. Cardio-myocytes grown on it were able to adhere andproliferate, indicating that TMC-DLLA co-polymersare not cytotoxic. In vitro signs of degradation occurredby 4 months via hydrolysis [103]. When these polymerswere implanted subcutaneously in rats, degradationstarted at 3 weeks and by one year 96% of its mass waslost. These co-polymers also elicited a sterile and acuteinammatory reaction, resulting in the formation of acapsule in vivo [104]. Degradation was far more rapid inpoly TMC polymers (3 weeks). Silica-based poly(di-methylsiloxane) (PDMS) has been polymerised withPLA-PGA composites to form PLGA/TEGOMERS,which have shown to be biocompatible, supportive tocell growth and non-toxic [105].Experiments with micro-vascular cell cultures mainly

involve PLA, PGA or their composite scaffolds. Theseare summarised in Table 4. PLs like poly CL arebiocompatible with prolonged degradation times of upto 3 years in vitro [106]. Combining CL with TMC slowsdown degradation of the primary TMC polymer [107].The mechanism of degradation is hydrolysis withcaproic acid, the by-product. Its disadvantage is that itis structurally weak.PPFs are polyesters based on fumarates. They can be

injected into any space and further photo cross-linkedby adding methacrylates [108]. PPFs are structurallystrong and capable of withstanding great loads. In vitrodegradation (up to 200 days) occurs following hydro-lysis of its ester bonds to form fumaric acid, a naturalmetabolite, propylene glycol and poly(acrylic acid-co-fumarate) [109]. Apart from being biodegradable, PPFhas also been found to be non-cytotoxic with minimalinammatory reaction in vivo [108].PA forms soft tissue constructs being biocompatible,

thermoplastic, with controlled degradation characteris-tics (surface erosion). It is a weak polymer limiting itsuse in load-bearing devices. It may be strengthened bythe addition of imide groups [110]. PAs are degraded byhydrolysis of its anhydride bonds by 12 months in vitrointo dicarboxylic acid [111]. In vivo implantation ofthese polymers shows evidence of good vascularisationof itself by 4 weeks [112].PCs are thermoplastic and strong enough to sustain

heavy compressive loads. Their advantages are that theyare biodegradable, biocompatible and have very lowrate of degradation. However, hydrolysis of its carbo-nate groups culminates in the formation of acidic by-products. This is an undesirable property for an ECMsubstitute [113].Bioelastic materials, monomeric protein polymers

PRESSals 26 (2005) 18571875 1865which convert heat or chemical energy into mechanicalenergy, represent the next generation of polymers. Heatexchange, hydrophobicity or hydrophilicity induces amechanical tensile force within the polymer that ismediated by competition between the molecules for

ARTIC

LEIN

PRESS

Table 4

Synthetic materials used for ECM construction in various tissues

Polymer Compressive

strength

Synthesis DT (months) By-products Advantages Disadvantages References

PGA 7.5GPa E/M/C 612 Glycolic acid Porous, biocompatible, non-

toxic, non-inammatory,

natural waste metabolite

Not pliable [99]

PLA 2.7GPa E/M/C 424 L-lactic acid Porous, biocompatible, non-toxic, slower degradation,

natural waste metabolite

Hydrophobic [113]

PLGA 1.9GPa E/M/C 1216 D,L-lactic acid Porous, biocompatible, non-

toxic, non-inammatory

Excessive leakage [100]

PL 0.4GPa E/M/C 424 Caproic acid Non-toxic, biocompatible Weak structure [106]PPF 230MPa Injection E30 Fumarate, polyethylene

glycol, poly(acrylic-co-

fumaric) acid

Non-toxic, biocompatible,

minimal inammation, photo-

cross-linkage, strong

Not signicant [108-109]

PA 1.3MPa Thermoplastic 12 (in vitro) Dicarboxylic acids Biocompatible, thermoplastic,

surface erosion, angioinductive

Weak [111]

PC Sufcient for load-

bearing

Thermoplastic Very slow Tyrosine, CO2 and alcohols Porous, biocompatible,

thermoplastic, osteoconductive

Acidic byproducts [181]

PU (LDI) 840Mpa Thermoset 12 Lysine, glycolic acid, caproic

acid

Biocompatible, strong, pliable Some PUs (TDI) are toxic [113]

PPZ Sufcient for load-

bearing

Phosphates and ammonia Biocompatible, osteoconductive,

strong

Not signicant [113]

POE Sufcient for load-

bearing

Regulatable

with lactide

Carboxylic acids Biocompatible, osteoconductive,

strong

Not signicant [182]

PEEU Tensile strain

(929MPa)

Biocompatible, strong Poor cell adhesion [183]

Ppy o13 Electrical conduction, minimalinammation

[184]

Keys: C, casting ; DT, degradation time; E, extrusion; M, moulding; , unknown.

R.Y

.K

an

na

net

al.

/B

iom

ateria

ls2

6(

20

05

)1

85

7

18

75

1866

hydration. These forces modulate cells within theconstruct to form micropatterns and secrete newECM. This process is called cellular tensegrity ormechanotransduction [114]. An example would bepoly(N-isopropylacrylamide), which can mechanicallyrecongure themselves with changes in hydrophilicityand temperature [98].

3.2.1.3. Solgel systems. Gels are composed of threecomponents namely solute, solution and voids, with thethree in equilibrium with one another [115]. A typical gelis polyethyleneglycol (PEG). As shown in Fig. 9,polymer molecules are initially functionalised withadhesion factors. These are then cross-linked via theformation of reactive termini. Cells migrating into this

ARTICLE INR.Y. Kannan et al. / Biomaterimatrix break down the cross-linking peptide leading tolocal degradation and hence release of the incorporatedfactors such as heparin and VEGF [95]. These syntheticgels need to be bio-degradable and cell-responsive [116].Using a PEG-based polymer, by shortening the blocklength from a mean molecular weight of 930 to6090 kDa, biodegradation of the hydrogel could beminimised [117]. The same effect has been shown intriblock co-polymers of poly(L-lactide)/PLLA andpoly(ethylene oxide) (PEO) (PLLA-PEO-PLLA) [115].Photo-polymerisable poly vinyl alcohol (PVA) hydro-

gels [118] are injectable. As such, it is possible to placethese gels in vivo using minimally invasive proceduresand then allowing cross-linkage to occur. This gel wasfound to be highly elastic and strong. Since celladherence was poor, these constructs were modiedwith the RGD peptide sequence [119]. This principle wasused by Cao and colleagues to immobilise ovalbuminwithin a 3-D hydrogel [120]. De Rosa and co-workersFig. 9. The synthesis of biomimetic gels.proposed a polyelectrolyte gel with a cationic surfacecharge across it. These charges increased the adhesion ofFb to the gel as well as their proliferation, therebyminimising the risk of brous encapsulation [121].

3.2.2. Superstructure

Superstructure refers to the integration of 2-D poresor structures into a 3-D conguration, whereby theseoverlap in a regular or irregular pattern. In polymers,the former confers crystallinity and the latter, amor-phous characteristics. The superstructure of a polymerdetermines pore characteristics. Studies have shown thatcells aggregate less within smaller pores and thesesmaller aggregates proliferate to a lesser extent [122].Hence, in scaffolds with smaller pore sizes and volumes,cells spread faster [123].Superstructures exist in three main forms namely

textile (woven/knitted), open-pore and angiopolar sys-tems. The angiopolar principle involves spatially or-ientating pores within matrices to induce cells,microvessels and nutrients to propagate in a givendirection [124]. Pores within biopolymers may beclassied into macropores (E500 mm), mesopores(2030 mm) and micropores (o10 mm). Macroporesregulate the optimum cell type and seeding densitywithin a construct. Mesopores allow microvessels toinltrate the polymer. It has been observed in early invivo studies on rats that arterioles and venules tend toproliferate on the inner surface of mesopores whilecapillaries lined its outer surface [124]. Lastly, micro-pores function as a regulator of molecular transportwithin a polymer system. Pore characteristics may bemodulated depending on the method of polymersynthesis like extrusion, leaching, casting, moulding orelectrospray.

3.2.3. Cells seeded onto matrices

Seeding cells onto a matrix is integral to the formationof a bio-hybrid scaffold as cells like broblasts andchondrocytes secrete ECM. However, long-term cellviability in such articial constructs is a limiting factoras these cells have to remain viable in vivo untilneovascularisation occurs. Fibroblasts are the mostabundant cells within the ECM, but scaffolds seededwith broblasts tend to contract following in vivoimplantation, thus impairing broblast migration. In astudy comparing pure and GAG-chitosan-collagensponges, the latter minimally contracted and showed a300% increase in broblast proliferation at 6 weekscompared to pure collagen sponges [125]. This wasaugmented by incorporating RGD-peptide sequences to

PRESSals 26 (2005) 18571875 1867improve broblast attachment and proliferation [126].Engineered cartilage using chondrocytes has the poten-tial to act as a viable ECM scaffold (as these cells havelow metabolic rates and require minimal perfusion), intowhich micro-vascular networks can actively proliferate

ated ex vivo [127] on PLGA scaffolds up to 2 months

further improves vascularisation prior to actual cell

INateriseeding [131].

3.3. Vascularisation

While scaffolds and microvessels are essential com-ponents of a capillary bed, it is equally important for thetissue-engineered device to connect to the host tissue.This would entail the induction of incoming micro-vessels into the scaffold (angioinduction) as well asallow the inherent microvessels to grow out and meetthese vessels (inosculation).

3.3.1. Inosculation

The presence of an inbuilt micro-vascular networkitself within a matrix is insufcient unless the intra- andperi-vascular regions are seeded with ECs, SMCs andbroblasts. Therefore, for the growth of vessels withinan engineered vascularised tissue construct out into thehost tissue, it is necessary for these to either link orhook-up with the hosts vasculature, this process is aterm described as inosculation. While ECs confer non-thrombogenicity to the graft, circulating haematopoieticstem cells such as EPCs are capable of stimulating theformation of a vascular network (post-natal vasculogen-esis) [134]. Their sources are the bone marrow, mono-following implantation, but viability decreased there-after. However, the study did not mention whether thesecells survived by neovascularisation or diffusion. Pre-adipocytes attached maximally to hyaluronic acidscaffolds with mean pore sizes of 400 mm [128]. Pre-laminating these cells layer onto layer in vitro improvedits vascularisation characteristics [129]. Fatty tissue fromlipoaspirate (LPA) has been shown differentiate intomany cell lines and hence form a source of stem cells[130]. Injecting the cDNA of VEGF into scaffoldsand grow. Embedded chondrocytes within brin glueshowed that these cells can maintain viability in a non-vascularised environment [132] although chondrocytesalone forms rigid tissues. Brown and co-workersdeveloped a hybrid chondrocyte-SMC with a compres-sive modulus between muscle and cartilage [133].Research has demonstrated that fat has potential as a

bulking agent; however, the difculties associated withthe extraction of adipocytes and their subsequent culture(the propensity for intra-cellular lipid vacuole bursting)has rendered their usage in tissue engineering to beextremely limited. Pre-adipocytes extracted from theepididymal fat pads of SpragueDawley rats prolifer-

ARTICLER.Y. Kannan et al. / Biom1868nuclear cells and the vessel wall. These cells can beharvested by enrichment following extraction or ex vivocultures [135]. Once these cells are injected into thecirculation, EPCs move on to sites of neovascularisation(therapeutic vasculogenesis) [136].Early studies on vasculogenesis used cells like quailblastodiscs, murine embryonic stem cells, HUVEC andhuman pulmonary micro-vascular endothelial cells(HPMEC) grown on collagen, brin, gelatin andmethylcellulose matrices (Table 5). These cells havebeen shown to grow into capillary-like tubes (CLTs),which may be controlled by altering the growth factorgradient or biomechanical tension across the gel. Whilein culture, tissues can be genetically modied bytransfecting them with viral vectors. Overexpressionand carcinogenicity are a problem. Alternatively, cellslike SMCs could be transduced with multiple genes andthen be allowed to repopulate the microenvironment(transgenesis). SMCs were harvested from tissue en-zymatically and cultured serially until differential celladhesion molecules are removed. Next, the SMCs weremagnetically extracted and incubated with high-titreretroviral supernatants. 99% of these cells were success-fully transfected [137].EC are sourced from arteries, veins, omentum and

subcutaneous fat [138]. Using enzymatic degradation,EC are harvested from vein grafts by either thecannulation or eversion technique and cultured for 56weeks. Grenier and co-workers have introduced amethod of obtaining EC, SMC and broblasts fromthe same vein biopsy [139]. Isolation of EC from fattytissue requires it to be minced, enzymatically digested,centrifuged in a Percoll gradient before extracting ECusing magnetic beads [138].Prior to in vivo implantation, EC or EPCs would need

to be seeded onto scaffolds. The seeding density woulddepend on the scaffold material, its porosity and themethod of cell inoculation. The standard mode forpromoting cell attachment is to culture cells onto thepolymer with or without adhesion molecules. Adisadvantage is that cells tend to adhere only to theouter layers of the scaffold. In order to achieve morehomogenous implantation throughout a 3-D scaffold,other methods like the drop on and low-pressurecentrifugation techniques which systematically seed cellsonto scaffolds have been introduced [128].Human adipose stromal cells placed within micro-

carriers and co-cultured with HUVECs were grown in aserum-free static culture model with the addition ofangiogenic factors. By day 16, stable CLTs with patentlumens were formed by outgrowing ECs. Larger loopscoalesced while smaller ones regressed, similar to thepruning phenomenon that is observed during embry-ological development of the vascular system. In vitro,this is termed guided migration.Clinically, the larger the microvessel, the greater the

PRESSals 26 (2005) 18571875perfusion [140]. An in vivo model wherein humandermal microvascular endothelial cells (HDMEC)seeded onto biodegradable polymer matrices were putinto immunodecient mice showed the following char-acteristics. On day 1 EC migrated through the matrix;

IN

indu

bryo

tokin

tokin

ateriARTICLE

Table 5

In vitro developmental models of angiogenesis and vasculogenesis

Cells CLS

formation

(day)

Morphogenesis

Bovine capillary EC 23 Spontaneous

Embryonic stem cells 12 Spontaneous em

body formation

Muscular tissue fragments adipose

tissue frag.

312 Spontaneous

Bovine aortic EC, adrenal capillary

EC, HUVEC

1 Spontaneous

Bovine capillary EC 515 Cytokines

Bovine capillary EC 23 Phorbol ester

Human umbilical vein EC, HDMEC 1 Spontaneous

Bovine capillary EC 12 Spontaneous cy

Rat aortic explants 7 Spontaneous

Bovine aortic EC 1018 Spontaneous

Human umbilical vein EC 1 Spontaneous

Rat fat microvessels fragments 46 Spontaneous

Calf pulmonary aortic EC 27 Spontaneous cy

Human umbilical vein EC, bovine

retinal EC

12 Spontaneous

R.Y. Kannan et al. / Biomon day 5 they formed CLTs and by 1 week haddifferentiated into functional microvessels. Maturationoccurred by 21 days and these microvessels eventuallylinked up with the host vessels[141,142]. Immunohisto-chemical analysis of a murine myocardial model showedthat monocytes/macrophages produced holes throughthe myocardium using proteolytic enzymes, the lumensof which were eventually colonised by EPCs [143].

3.3.2. Angioinduction

The process of inducing incoming microvessel forma-tion depends on the growth factors used and the methodof delivery. Their angiogenic effects were studied usingin vivo models like chorio-allantoic membrane (CAM)[144,145]. FGF and VEGF are mainly involved invasculogenesis. Angiogenic sprouting is mediated bytransforming growth factor-b (TGF-b), while matura-tion of vessels is via angiopoietin-1 and -2 (Ang-tie) andplatelet-derived growth factor (PDGF) pathways [146].Their effects are summarised in Table 6.The primary stimulus and factor causative for the

process of capillary sprouting is hypoxia.Within 30min of its onset, early growth response

factors (Egr-1 and -3) are expressed [147]. VEGFmodies the ECM to allow EC and broblast migration[140,30], acting on the venules, it increases their

Embryonic stem cells 11 Spontaneous embryo

body formation

Human placental blood vessels 721 Spontaneous

Mice microvessel fat pad 14 Spontaneous cytokin

Bovine aortic EC, human HUVEC 3 Spontaneous

Human marrow microvascular EC 2150 Spontaneous cytokin

Keys: EC, endothelial cells; HUVEC, human umbilical vein endothelial

bronectin.PRESS

ction Matrix Spatial

organisation

References

Collagen I sandwich 3-D [156]

id Culture dish 3-D [157]

Fibrin+Collagen I 3-D [158]

Fibrin 2-D [159]

Collagen I 3-D [160]

Fibrin 3-D [161]

Matrigel 2-D [162]

es Fn+Collagen

IV+Matrigel

2-D [163]



Fibrin I or II sandwich 3-D [166]


es Fibrin in micro-carriers 3-D [168]

Fibrin 2-D [169]

als 26 (2005) 18571875 1869permeability. Subtypes A and B are generally involvedin EC migration and proliferation, while C and Dsubtypes are responsible for venous and lymphaticproliferation [25]. FGF is another angiogenic factorused. Derived from SMCs and ECs, FGF is stimulatedby EC regeneration, hypoxia and collateral formation toelicit cell induction and proliferation. In pre-fabricatedaps, Bayati and colleagues showed that tissue viabilitywas enhanced by using FGF [148].Monocyte chemoattracted protein-1 (MCP-1) is

released following shear stress to the vessel wall. Itattracts circulating and in-situ MCs to home in at sites ofneovascularisation [149]. Granulocyte-Macrophage Col-ony Stimulating Factor (GM-CSF) acts synergisticallywith MCP-1 by promoting arteriogenesis, stimulatingMC/MP release from the bone marrow as well asprolonging their life span [17]. TGF-b described earlieralso serves as a growth factor which, depending on thereceptor, either activin-like kinase type-1 (ALK-1) orALK-5, stimulates or inhibits angiogenesis, respectively[150]. This illustrates the overlapping functions of thesemediators. The slow release of these growth factorswould diminish function, while excessive amountswould induce brovascular growth.Control over spatial and temporal patterns of the

chemicals release is necessary to form an organised,

id methylcellulose 3-D [170]

Fibrin 3-D [171]

es Collagen gel 3-D [185]


es Fn+Collagen I 2-D [187]

cells; HDMEC, human dermal microvascular endothelial cells; Fn,

bioactive scaffolds [152], preferably surface-erodingscaffolds like poly(glycerol sebacate) (PGS) [153].Other

IN

Recep

FGFr

VEGFR2/Flk-1 EC formation

VEGF

VEGF

ALK-

Tie-1

PDGF

Tie-1

Eph B

wth f

aterimethods of delivery include viral transduction [154],direct inoculation [155] or genetically modied cells[137].

4. Conclusion

The biggest hurdle in the construction of tissue-engineered aps is the inability of existing synthetic ortissue-engineered small-diameter vascular grafts(o6mm) to sustain ow through them and, moreimportantly, the absence of any source of an incorpora-congured and functional micro-vascular network. Invitro, growth factors may be incorporated by co-culturing them with constructs or placing these factorswithin bioarticial organs like microcarriers which allowregulated release of chemicals. In vivo, these factors areintroduced directly into the bloodstream, incorporatedinto cells or packed into defects as beads [151] or

ARTICLE

Table 6

Role of angiogenic growth factors during vascular development [146]

Event Factor

Angioblast induction FGF

Conversion of angioblasts to EC VEGF

EC formation into tubes VEGF

Angiogenic sprouting VEGF

EC activation TGF-b1Vessel thickening Angiopoietin-1

Smooth muscle recruitment PDGF-bVessel thinning Angiopoietin-2

Arterio-venous differentiation Ephrin B2

Keys: FGF, broblast growth factor; VEGF, vascular endothelial gro

derived growth factor-beta.

R.Y. Kannan et al. / Biom1870table capillary bed, natural or otherwise. Current trialson EC-seeded infra-popliteal vessel bypass grafts haveonly been proven to be as successful as vein grafts in thesame setting [32], while the development of an articialcapillary bed is limited to in vitro models [156171].However, a better understanding of microuidics andfurther breakthroughs in nanotechnology may makethis more likely.

4.1. Future perspectives

Micro-contact printing, printing biological moleculesdirectly onto scaffolds, is spawning a new generation oftissue-engineered constructs. Here, a blueprint is madeusing a computer-aided design and the construct is thenbuilt within 24 h. For instance, gold or silver plates areinitially patterned with alkane-thiolates onto whichbronectin molecules are attached. In turn, bronectinbinds to EC in culture [172]. These individual celldroplets then coalesce due to tissue uidity. Whenapplied in many layers, three-dimensional constructscan then be developed. This technique has been used toform articial liver micro-channels [173]. As the keyissue is tissue or cell perfusion, these constructs arenally placed within a bio-reactor [174]. Recently, aporous endotheliased network based on human pul-monary endothelial cells was successfully developedusing this technology [175]. Once popularised, thistechnique would revolutionise fabrication of precisetissue-engineered constructs.Perhaps, the most exciting tissue engineering advance

is nanotechnology. Earlier generations of biomaterialshad micro-dimensions while biomolecules are nanos-tructures. As such, physiological processes within thebody could not be modulated by these devices until now.The recent discovery by Moldovan and colleagues thatECs and EPCs in the bloodstream are capable ofliterally drilling holes through matrices has opened up anew perspective. These groups have proposed buildingan angiogenesis assist device or angiochip usingnanotechnology, wherein foci of EC would be spatially

R2/Flk-1 Primitive vascular plexus

R-2/Flk-1, VEGFR-1/Flt-1 Angiogenesis

1, ALK-5 Angiogenic modulation

Vessel stabilisation

R-b Vessel stabilisationand Tie-2 Antagonistic to angiopoietin-1

4 Remodelling

actor; TGF-b1, transforming growth factor beta-1; PDGF-b, platelet-PRESS

tor Action

Angioblast formation

als 26 (2005) 18571875distributed within a three-dimensional matrix intercon-nected by articially created grooves for proliferationand subsequent tube formation [172]. This device iscurrently being used to develop liver tissue. Similarly,they could be spatially distributed within any tissue andact as angioconductive and angioinductive foci. Ulti-mately, it is the temporal synergism achieved betweenvessel, matrix and cell, which would determine thecreation of a vascular blueprint for tissue engineering.

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[8] Seifalian AM, Tiwari A, Hamilton G, Salacinski HJ. Improving

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[9] Tiwari A, Salacinski HJ, Punshon G, Hamilton G, Seifalian

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