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Journal of Biomaterials Science, Polymer Edition

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Experimental approaches to vascularisation withintissue engineering constructs

Md. Sarker, X.B. Chen & D.J. Schreyer

To cite this article: Md. Sarker, X.B. Chen & D.J. Schreyer (2015) Experimental approaches to

vascularisation within tissue engineering constructs, Journal of Biomaterials Science, PolymerEdition, 26:12, 683-734, DOI: 10.1080/09205063.2015.1059018

To link to this article: http://dx.doi.org/10.1080/09205063.2015.1059018

Accepted author version posted online: 08 Jun 2015.

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REVIEW ARTICLE

Experimental approaches to vascularisation within tissue engineeringconstructs

Md. Sarker a *, X.B. Chena ,c and D.J. Schreyer a , b

a Division of Biomedical Engineering, College of Engineering, University of Saskatchewan, 1A26,57 Campus Drive, Saskatoon, SK, S7N 5A9, Canada; b Department of Anatomy and Cell Biology,College of Medicine, University of Saskatchewan, 5800 Saskatoon City Hospital, Saskatoon, SK,S7K 0M7, Canada; c Department of Mechanical Engineering, College of Engineering, University

of Saskatchewan, 3B47, 57 Campus Drive, Saskatoon, SK, S7N 5A9, Canada

( Received 26 March 2015; accepted 2 June 2015)

Tissue engineering opens up a new area to restore the function of damaged tissue or replace a defective organ. Common strategies in tissue engineering to repair andform new tissue containing a functional vascular network include the use of cells,growth factors, extracellular matrix proteins, and biophysical stimuli. Yet, formationof well-distributed, interconnected, and stable vascular network still remainschallenging. In addition, anastomoses with host vasculature upon implantation andlong-time survival of the new blood vessel in vivo are other critical issues to beaddressed. This paper presents a brief review of recent advances in vascularizationin vitro as well as in vivo for tissue engineering, along with suggestions for futureresearch.

Keywords: vascularization; angiogenesis; vasculogenesis; biomaterial; growthfactors; scaffold

1. Introduction

A major challenge in the creation of new tissues by tissue engineering is the limitation

of diffusional mass transfer into new tissue having a thickness of more than

100 – 200 μm,[1] causing ischemia and cell death. Vascularization overcomes the limit

of diffusional mass transfer for delivery of various types of nutrients, oxygen, growth

factors, biochemical signaling factors, and removal of carbon dioxide, and metabolic

waste.[2] The 100 – 200 μm limit regulates the critical distance necessary between two

successive capillaries to prevent an ischemic condition.[3] This issue must beconsidered in any experimental or therapeutic effort to regenerate functional tissue of

appreciable size using a tissue engineering approach.

Two mechanisms are responsible for the generation of a vascular network in vivo:

vasculogenesis and angiogenesis. Vasculogenesis generally takes place at an embryonic

stage, while angiogenesis occurs mostly in postnatal life.[4,5] The term vasculogenesis

denotes the assembly of endothelial progenitor cells (EPCs) that eventually form capil-

laries and a vascular network [6] in response to various growth factors, then undergoes

remodeling to form functional vasculature. Angiogenesis refers to the sprouting of new

blood vessels from existing ones due to specic angiogenic signals.[4] Engineered

*Corresponding author. Email: [email protected]

© 2015 Taylor & Francis

Journal of Biomaterials Science, Polymer Edition, 2015

Vol. 26, No. 12, 683 – 734, http://dx.doi.org/10.1080/09205063.2015.1059018

mailto:[email protected]://dx.doi.org/10.1080/09205063.2015.1059018http://dx.doi.org/10.1080/09205063.2015.1059018mailto:[email protected]


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tissue can be vascularized by either process, but must ultimately result in anastomosis

with host vasculature.

2. Different types of blood vessels and their functions

The vascular system is a geometrically complex network consisting of arteries that

supply blood to tissues, and veins that remove it (Figure 1). Major arteries branch out

into successively smaller arterioles, metarterioles, and arterial capillaries. Capillaries

combine downstream to form venules, and then major veins.

Arterioles typically have diameters in the range of 10 – 300 μm and are composed of

concentric layers of smooth muscle cells (SMC), basement membrane and endothelial

cells (ECs).[7] A metarteriole is a short vessel which branches out from an arteriole

with a smaller diameter. Generally, a metarteriole consists of only two layers; SMCs

surrounding ECs. At the point where capillaries branch out from metarterioles, there

exists a loop of SMCs called the precapillary sphincter. The function of precapillary

sphincter is to control the blood ow to the capillary bed and to ensure enough time tofacilitate diffusion and other mass transfer mechanism between capillaries and nearby

cells. To keep the blood ow continuous, a bypass vessel called the thoroughfare chan-

nel connects the metarteriole with the venule. The capillaries consist only of ECs hav-

ing an outer layer of protein-based basement membrane. However, pericyte cells, which

are responsible for paracrine signaling and capillary blood ow, can be found within

the capillary basement membrane. The capillaries are also known as microvessels, and

have a diameter of 4 – 10 μm. The transition from capillary to venule takes place gradu-

ally and the approximate diameter of post-capillary venule is 10 – 50 μm. Post-capillary

venules then fuse with a connecting venule whose diameter is 50 – 300 μm.[8]

Capillaries are the principle site of mass transfer. The difference between the pressure of blood within the capillaries and the interstitial uid pressure causes

the movement of uid into or out of capillaries. Generally, three types of capillaries are

Figure 1. Blood circulatory system inside the body (Reproduced from [288]).

684 Md. Sarker et al.


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seen in the body: continuous, fenestrated and sinusoidal (Figure 2). Most capillarieswithin the body are the continuous type. These have tight junctions as well as inter-

cellular clefts between the ECs, and are covered with basement membrane. Water, gas

molecules, ions, and other water soluble molecules can pass in or out of continuous

capillaries through diffusion, vesicles, and intercellular clefts. However, high molecular

weight molecules are transported across ECs by pinocytosis. Fenestrated capillaries

have 60 – 80 nm pores in their ECs and are covered with basement membrane.

Fenestrated capillaries carry out mass transfer through diffusion, vesicles, intercellular

clefts, and pores. This type of capillary is found within the endocrine glands, intestines,

pancreas, and kidneys. Sinusoidal capillaries are leaky in nature and have 30 – 40 μm

pores, much larger than those seen in fenestrated capillaries. Sinusoidal capillaries have

incomplete basement membrane covering and more intercellular clefts compare to other

capillaries. White and red blood cell and serum protein can pass through the walls of

sinusoidal capillaries. Generally, sinusoidal capillaries are found in bone marrow, liver,

spleen, and adrenal gland.[9]

3. Mechanisms of blood vessel formation in vivo

3.1. Vasculogenesis

EPCs from bone marrow or blood are mainly responsible for pre- and post-natal vascu-

logenesis through a complex series of steps (Figure 3). At the initial stage, EPCs move

from their source and enter unvascularized tissue in response to chemo-attractant fac-

tors, such as granulocyte – monocyte colony stimulating factor, granulocyte colony

stimulating factor, vascular endothelial growth factor (VEGF), basic broblast growth

factor (bFGF), placental growth factor, erythropoietin and stromal cell-derived factor-1

(SDF-1).[10] Moreover, factors secreted from ischemic tissue cause nitric oxide (NO)

production that activates extracellular proteases, especially matrix metalloproteases-9

(MMP-9), and facilitates EPC migration.[11]

Homing of EPCs into ischemic tissue site requires chemotaxis, adhesion, and trans-

endothelial migration. After moving out from bone marrow traveling through blood

vessels, EPCs undergo chemotaxis toward a particular ischemic region in response to

gradients of chemokines, such as SDF-1, interleukin-8 (IL-8), growth-regulated onco-gene-α, and C – C chemokine.[12,13] The chemokines also stimulate EPCs to promote

adhesion with the inner layer of blood vessels. EPCs pass through the endothelial lining

Figure 2. Structure of different types of capillaries (Reproduced from [288]).

Journal of Biomaterials Science, Polymer Edition 685


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of blood vessels by a specic transendothelial migration mechanism. Once at the base-

ment membrane, EPCs initiate their invasion mechanism by rupturing it and remodeling

basement membrane proteins, as well as the extracellular matrix (ECM) of the intersti-

tial space, using extracellular proteases, such as MMP-9, cathepsin L, urokinase-type

plasminogen activator, and tissue-type plasminogen activator, to reach the ischemic tis-

sue site.[14,15] Then EPCs accumulate at the ischemic site; form a vascular pattern bydifferentiating into ECs and interacting with existing ECs and ECM. EPCs attached to

the ECM proliferate under the inuence of VEGF, immunoglobulin, and epidermal

growth factor (EGF),[16,17] and differentiate due to monocyte chemoattractant protein-

1, insulin-like growth factor-1 (IGF-1), SDF-1, VEGF, and platelet-derived growth

factor (PDGF).[18,19]

3.2. Angiogenesis

During angiogenesis, new capillaries sprout from existing blood vessels in response to

various biochemical signals and expand into ischemic tissue following tracks, gradients,or attractive or repulsive signals. Any unnecessary vascular plexus caused by random

sprouting undergoes remodeling according to the requirements of specic tissue.

Figure 3. Blood vessel formation by vasculogenesis (Reproduced from [289]).



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Functional non-uniformity of ECs along the blood vessel wall caused by VEGF

gradients plays a critical role in angiogenic sprouting. Mechanistically, specic ECs

respond to a chemo attractant by extending as tip cell, then propagating in different

directions by forming stalk-like structure, which eventually acquire a lumen.[20,21] To

communicate with the surroundings, tip cells extend multiple lopoida which are sensi-

tive to various guidance molecules and specic growth factors, such as VEGF-A, FGF,

angiopoietin-1 (Ang-1), EGF, and IL-8. The spatial as well as temporal distributions of

these factors are very crucial for guiding the growth of new capillaries within the

ischemic tissue.[22] For example, a VEGF-A gradient in tissue plays an important role

for lopodia extension from the tip cell, and the corresponding receptor for VEGF-A is

VEGFR-2, which is expressed by the tip cell.[23] Angiogenesis in ischemic or newly

formed tissue is a complicated process which can be demonstrated through some steps

(Figure 4).

At the beginning, ischemic tissue secretes VEGF which dilates the existing blood

vessel through the redistribution of intercellular adhesion molecules as well as mod-

ication of the cell membrane structure. It also increases vascular permeability with thehelp of NO [24,25] and assists the relocation of plasma protein to form temporary

structures to facilitate EC migration. Besides VEGF, Ang-2 contributes to angiogenic

sprouting through the degradation of ECM protein and the removal of SMCs from the

outer surfaces of capillaries.[25,26] Degradation of basement membrane and ECM are

vital for EC migration, while microvascular ECs themselves secrete some matrix met-

alloproteases, such as MMP-2, MMP-3, MMP-9, along with protease inhibitor mainly

tissue inhibitor, of metalloproteinase-2, which coordinately regulate basement mem-

brane and ECM degradation [27] by causing various spatial and temporal factors to

Figure 4. Blood vessel formation by angiogenesis in brain cancer (Reproduced from [290]).



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generate non-uniformity along the wall of blood vessel.[28,29] In addition, another

important feature of extracellular proteolytic degradation is the release of some growth

factors, such as IGF-1, VEGF, and FGF-2, which also facilitate angiogenic

sprouting.[30]

The progression step is associated with the migration and proliferation of ECs due

to the inuence of VEGF, angiopoietins (Angs), FGFs, and PDGF. Among them, Ang-

1 regulates the interaction between ECs and periendothelial cells during migration,

[25,31] FGF facilitates EC growth via the recruitment of mesenchymal cells,[ 32] and

PDGF enhances EC growth through the recruitment of pericytes and SMCs.[33] In

addition, other factors involved in EC proliferation and migration include neuropep-

tides, IGF-1, erythropoietin, hepatocyte, and interleukins.[34,35]

After migration and proliferation of EC into ECM, ECs eventually differentiate to

form cord-like structure and a lumen starts developing due to the inuence of VEGF

and Ang. Interestingly, Ang-1, VEGF121, and VEGF165 help to increase lumen

diameter, while VEGF189 decreases lumen diameter.[36]

4. Tissue engineering approaches to blood vessel formation

Well-distributed and interconnected vascular network is required to maintain the viabil-

ity of large cell population in engineered tissue. While tissue engineering is eventually

emerging as a potential approach to regenerate tissue and organ, formation of 3D

vascular network within the neotissue still remains challenging. Current strategies to

form complex vascular channel within engineering construct or growing tissue can be

categorized as scaffold-based and scaffold-free approaches (Figure 5). Strategies

(Table 1) associated with scaffold-based approaches are, construction of 3D matrix with

appropriate biomaterial by biofabrication technique or stacking multiple micropat-terened thin planar surfaces or micro tissue modules, engineering new tissue with

decellularized matrix, functionalization of scaffolds with cells, angiogenic factors (AFs)

or proteins, transfection of seeded cells by gene therapy to obtain sustained release of

AFs and prevascularisation to maintain viability of the seeded cells. In contrast, in scaf-

fold-free approach cell sheets are assembled to promote vascularized tissue.

4.1. Scaffold-based aproaches

4.1.1. Scaffold fabrication biomaterial

The choice of matrix biomaterial has profound impact on scaffold vascularization in tis-

sue engineering. Like natural ECM, the biomaterial should facilitate vascular cell –

biomaterial interactions, which would activate numerous signaling pathways to promote

vascular cell survival, proliferation, differentiation, and migration. A tissue scaffold

experiences different types of force upon implantation in vivo, such as compression,

shear, torsion, and tensile force. Therefore, the biomaterial should have suf cient

mechanical strength to sustain those forces to keep the scaffold architecture intact.

Moreover, controlled biodegradability is desired so that over time, the scaffold could

maintain its mechanical strength while the biomaterial is replaced with a vascular net-

work and newly formed tissue. Furthermore, cytotoxicity has to be considered for its

impact on implanted or host cell survival in vivo or during preparation in vitro.

Generally, polymers are used as biomaterials for the fabrication of scaffolds.[37]Biopolymers can be classied into three categories in terms of preparation or source:

synthetic, natural, and hybrid.



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4.1.1.1. Synthetic polymer. Synthetic polymers used for tissue engineering offer

scaffold fabrication exibility while maintaining desired mechanical properties of the

engineering construct. However, lack of cell adhesion peptides in the molecular struc-

ture of synthetic materials limits their application for vascular tissue engineering. A

number of biodegradable, biocompatible, and non-toxic synthetic polymers have shown

potential for tissue engineering applications. Among them, polyglycolic acid (PGA),

polylactic acid (PLA), polylactic-co-glycolic acid (PLGA), poly-L-lactic acid (PLLA),

poly-ε-caprolactone (PCL), polyethylene glycol (PEG), and polyhydroxyalkanoates

(PHAs) have frequently been explored in different studies of tissue-engineered graft

vascularization.[38]

When exposed to pulsatile conditions, PGA mesh tube scaffolds seeded with

porcine carotid SMCs were eventually replaced by blood vessel-like structures while

maintaining uniform SMCs distribution and collagen content close to that of native ves-

sels.[39] In vivo experiments have also shown PGA scaffolds as promoters of vascular-

ization. PGA scaffolds seeded with bovine artery cells degraded completely after

11 weeks of implantation in the pulmonary artery of sheep and were replaced by new

tissue having 73.9% collagen content with respect to native tissue.[40] Similarly, in an

in vivo study, PEG diacrylate hydrogel was implantated subfascially into Sprague-

Dawley rats and kept in vivo for three weeks. Interestingly, even in the absence of

incorporated growth factors, vascularized tissue ingrowth was reported throughout the

entire PEG gels having pore size 50 – 150 μm.[41]Derivatives of PLA (polymer of lactide) have also shown their angiogenic potential

in different studies. For example , in vivo implantation of tubular scaffold fabricated

Figure 5. Different approaches for vascular network formation.



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with porous lms of poly-D, L-lactic-co-glycolic acid facilitated the ingrowth of

brovascular tissue, which eventually formed a vascularized, tubular tissue.[42] To

investigate the combined effect of PLLA and collagen gel on vascularization, porous

PLLA scaffold was exposed to in vitro condition, where the pores were lled with the

mixture of human aortic SMCs and collagen gel, resulting in the rapid formation of a

smooth inner layer around the vascular graft.[43] Another composite polymer of lactidePLGA was explored to study the release prole of VEGF. VEGF was directly incorpo-

rated into PLGA scaffolds and pre-encapsulated in PLGA microspheres. It was reported

that direct incorporation of VEGF resulted in 40 – 60% release within ve days in vitro,

while encapsulation of VEGF within microspheres resulted in prolonged release. When

implanted into subcutaneous pockets of SCID mice, VEGF released from PLGA

scaffolds signicantly enhanced local angiogenesis.[44]

Like PLA, composites and derivatives of PCL have been identied as a vasculature

enhancer in several studies. In a comprehensive review, Woodruff et al. presented the

attractive features of PCL (e.g. inexpensive, FDA-approved, and superior rheological

and viscoelastic properties) to demonstrate its applicability for the fabrication of longer

term degradable implants suitable for a specic anatomical site.[45] In vivo implanta-

tion of PCL/PLA tubular scaffolds seeded with mixed cells obtained from femoral

Table 1. Tissue engineering approaches to blood vessel formation.

Approaches Description

A. Scaffold based approachA.1. Scaffold fabrication

biomaterial

Synthetic polymers (e.g. PGA, PLA, PLGA, PLLA, PCL, PEG,

and PHAs); natural polymers (e.g. collagen, bronectin, brin,elastin, silk broin, HA, alginate, chitosan); decellularized matrix(e.g. allogenic or xenogenic tissue or organ graft)

A.2. 3D printingapproaches

3D Fabrication (e.g. 3D plotting, inkjet printing, laser-based biofabrication and stereolithography) technique to fabricatecomplex vascular network; microfabrication (e.g. microuidics,micromolding, lithographic, direct laser write) technique to indent micro pattern on polymer or cell seeded hydrogel to allowcomplex network formation at micro level; modular assembly of

patterned micro modules generated from self-assembledaggregation, microfabrication of cell seeded hydrogels, or direct

printing method to form macro tissue assembled by random

packing, stacking of layers or directed assemblyA.3. Addition of cells Autologus mature ECs (e.g. human umbilical vein, human dermalmicrovascular, human vascular), stem cells (e.g. embryonic stemcell, MSC), EPCs (e.g. peripheral blood derived, umbilical cord

blood derived), vascular wall cells (e.g. pericytes, vascular SMCs, pulmonary artery SMCs)

A.4. Addition of growthfactors

VEGF, FGFs, PDGF, Angs, TGF-β

A.5. Gene therapy Viral vectors (e.g. retrovirus, lentivirus, adeno-virus and adeno-associated virus) and nonviral vectors (e.g. plasmid DNA)

A.6. Prevascularisation In vitro or in vivo culture prior to implant to form vasculature andmaintain cell viability

B. Scaffold free approachB.1. Cell sheet technique Temperature controlled cell attachment, culture and detachment

followed by stacking multiple layers



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veins of a mongrel dog showed no evidence of stenosis, thrombus, occlusion, or

aneurysmal formation within the tubular graft. Moreover, the vascular graft was

replaced by native tissue containing a large amount of ECM, where the ECs were

found to organize themselves in a linear way on the luminal surface of scaffold.[46]

Besides animal model study, in May 2000 a clinical trial was done on a four-year-old

girl to implant a pulmonary bypass graft made of PCL/PLA copolymer seeded with

autologous vein cells. This implantation was a milestone in pediatric cardiovascular sur-

gery as the patient was doing well even after seven months of implantation and no

postoperative complications were reported.[47] Another derivative of PCL called poly

(lactide-co-caprolactone) (PLCL) was studied for tissue vascularization due to its bio-

compatibility, elasticity, and slow degradation property. In a study, PLCL vascular

grafts seeded with rabbit aortic SMCs were exposed to pulsatile ow in a perfusion

bioreactor for eight weeks. It was reported that the PLCL graft signicantly regulated

the proliferation of SMCs, increased collagen production, and promoted cell alignment

similar to that of native vascular smooth muscle tissues.[48]

The angiogenic potential of PHA derivatives and its composites were demonstratedin different investigations. In a study, PHA – PGA tubular scaffolds seeded with ovine

carotid artery cells were implanted into abdominal aortic segments of sheep for ve

months. No aneurysms were observed in the tubular graft; however, the percentage of

collagen, DNA content, as well as the mechanical strength of grown tissue, reached

close to that of native vessels.[49] Another research group maintained static/pulsatile

condition for PGA/PHA derivative-fabricated vascular grafts seeded with ovine vascular

myo broblasts and ECs. After 28 days of culture, it was reported that all samples

gained viable, dense, and conuent smooth tissue with high collagen content.[50]

4.1.1.2. Natural polymer. Biopolymer or naturally derived polymers can be classiedinto two categories, such as; protein-based polymers and polysaccharidic polymers. In

tissue engineering, natural polymers are more advantageous over synthetic polymers in

many aspects, however, are not free from shortcomings.Some advantages and

drawbacks of natural polymers have been presented in Table 2.

Protein-based polymers, particularly collagen, brin, bronectin, elastin, and silk

broin, have frequently been explored in different studies for vascularization applica-

tion. For example, it has been reported that freeze-dried collagen scaffold seeded with

dermal broblasts and bone marrow-derived stem cells enhance the proliferation of

EC.[51] Further, immobilized VEGF on collagen scaffolds promotes EC invasion and

proliferation.[52] Another group incorporated modied VEGF into collagen scaffolds

implanted subcutaneously into mice, resulting in an improved vascular network within

the implanted scaffold.[53] Besides, Fibrin tissue engineering scaffolds have also pre-

sented promising results for promoting blood vessel formation. Human microvascular

ECs were seeded into a 90% brin, 10% collagen matrix with combinations of added

growth factors; FGF-2 or VEGF, with TNF-α.[54] This investigation suggests that b-

rin matrix facilitates the formation of a vascular capillary network through matrix bri-

nolysis by ECs, which occurs from the activities of plasmin and metalloproteinases. In

addition to providing specic cell adhesion sequences, a brin scaffold can store and

release angiogenic growth factors, such as VEGF-A165 and FGF-2, as well as plasmid

DNA to improve angiogenesis.[55] Another ECM protein bronectin has been consid-

ered as a vasculature promoter due to its enormous contribution in vascularization. In astudy, bronectin-coated collagen modules increased human vascular EC survival and

vessel formation when implanted in immunodecient mice compared to collagen



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Table 2. Natural biomaterials for scaffold fabrication.

Biomaterial Advantages Disadvanages

Collagen (1) Biocompatible, biodegradable, ease

availability and modiability,nonantigenic and controlled

protein release ability.[291](2) Hydrophilic and enhance cell

interactions.[292]

(1) Poor mechanical strength,[293] deformability,

exibility and tensilestrength.[294]

(2) High cost of purication,handling complexity,[295]rapid degradation, littlecompression forceresistivity.[296]

(3) Sterilization causes alterationof protein structure.[291]

Hyaluronicacid

(1) Easily modiable,hydrophilic, nonadhesive,

biodegradable,[297]nonantigenic and non-inammatory.[298]

(2) Regulate cellular migration,interaction, anddifferentiation.[299]

(1) Requires purication to avoiddisease transmission.[300]

(2) Rapid degradation rate, poor mechanical strength,[301]and forms scar tissuein vivo.[294]

Fibrin (1) Non toxic, non inammatory,[302] non allergenic,[303]non immunogenic[304],economical,[305]

biodegradable [306] and

easily processable.[307](2) Facilitates synthesis of collagen [308] and promotesangiogenesis.[309]

(3) Enhances attachment,migration and proliferation of smooth muscle [310] andECs.[311]

(1) Inconsistency in polymerization, fragile andstress squeezable.

(2) Structural weakness.[312](3) Upon implantation, solubility

increases and structuralintegrity decreases over time.[313]

Fibronectin (1) Critical role for angiogenesis[314] and scaffoldvascularisation.[315]

(2) Facilitates cell adhesion,[316]migration,[317] proliferationand alignment.[318]

(3) Can be used to coat articial biomaterials to support celladhesion, spreading [319] and

proliferation.[320]

(1) High cost.[321](2) Excessive mechanical stress

interrupts formation of

vasculature.[322]

Alginate (1) Non inammatory,[321] biocompatible, non toxic,cheap and potential for cellencapsulation and drugdelivery.[323,324]

(1) Uncontrollable degradationkinetics due to loss of divalent ion and burst releaseof protein drugs at higher

pH.[325]

(Continued )



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modules without bronectin coating.[56] Similarly, implantation of a bronectin/

collagen I gel seeded with EC and mesenchymal cell facilitated the formation of well-

perfused and long-lasting vascular network after implantation into mice.[57]

The role of structural protein elastin in vascularization is very signicant as geneti-cally modied mice lacking elastin died within four days of birth from arterial obstruc-

tion resulting from subendothelial cell proliferation and reorganization of smooth

muscle into the lumen of artery.[58] Cross-linked soluble alpha-elastin discs were found

to facilitate vascular SMC adhesion,[59] whereas tropoelastin-coated surface was

reported as a promoter of EC attachment and proliferation.[60] Elastin composites have

also been investigated to vascularize tissue engineering grafts. For instance, in a study,

human coronary artery SMCs were cultured for seven days on aligned nano brous

polyurethane scaffolds blended with the mixture of elastin and collagen. It was con-

cluded that the growth of SMCs on elastin/collagen-blended scaffold increased by

224% compared to that of polyurethane.[61] Besides animal body-derived protein, silk-

worm-secreted protein silk broin showed promising results in several studies in tissue

vascularization. In a study, co-culture of human dermal microvascular endothelial cells

(HDMECs) and primary human osteoblast cells (HOCs) in 3D silk broin nets caused

ECs to form intertwined microcapillary-like structures.[62] In another study, 3D silk

broin scaffolds pre-cultured with HDMEC and HOC were implanted into

immunodecient mice. After 14 days, it was found that the microcapillary structures

pre-formed in vitro, perfused well with the host vasculature.[63]

Apart from protein-based polymers, polysaccharide polymers, mostly hyaluronic

acid (HA), alginate, and chitosan have been studied for tissue vascularization. For

example, subcutaneously injected HA combined with recombinant gelatin facilitates the

formation of vascular network as well as the deposition of ECM components inrats.[64] Another group achieved controlled release of two angiogenic growth factors

(VEGF165 and PDGF-BB) from a hybrid mesh consisting of poly (3-caprolactone)-

collagen blend and HA hydrogel. They reported that the growth factor-loaded hybrid

mesh enhances cellular attachment and formation of vascular capillary network within

the tissue engineering construct during the culture of ECs and broblasts in a 3D

model.[65] However, HA has limitations in tissue engineering applications as it exhibits

poor cell attachment and mechanical properties.

Seaweed-derived polysaccharide alginate does not have EC adhesion properties, but

can be introduced through adsorption of proteins,[66] or by covalent incorporation of

different functional groups into the sugar backbone.[67] To facilitate vascularizationwithin alginate scaffolds, internal pores should be well interconnected and large enough

to allow ingrowth of blood vessels upon implantation in vivo.[68] For example, an

Table 2. (Continued ).

Biomaterial Advantages Disadvanages

(2) pH sensitive behavior facilitates drug delivery.[325]

(2) Poor mechanicalstrength.[326]

(3) Hydrophilic character prevents it from adsorptionof cell adhesive protein.[327]

(4) Lack of cell adhesion.[323]



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alginate scaffold with 90% porosity and pore sizes ranging from 50 to 200 mm

facilitates human embryonic stem cell aggregation and the formation of void and tube-

like structures in vivo.[69] Furthermore, growth factors can be incorporated in alginate

scaffold to enhance vascularization, for instance incorporation of VEGF in alginate

hydrogels enhance neovascularization into the matrix.[70,71]

Due to remarkable wound healing capability, chitosan and its composites were

explored for tissue engineering graft vascularization in different studies.[72] In a study,

porous chitosan, glycosaminoglycans (GAGs)-chitosan and dextran sulfate (DS)-chi-

tosan scaffolds were implanted in dorsal subcutaneous pockets of male Sprague-Dawley

rats. The rats were sacriced after two weeks and it was noticed that chitosan, heparin –

chitosan, and DS-chitosan scaffolds promoted cell proliferation, tissue ingrowth, and

formation of vascularized granulation tissue. Chitosan composites can also be used for

the controlled release of growth factors. To study the release prole of an AF from chi-

tosan complexes, FGF-2 incorporated chitosan/heparin hydrogel was subcutaneously

injected into the back of male Sprague-Dawley rats. Controlled release of FGF-2 mole-

cules for 20 days caused micro blood vessel and brous tissue formation around theinjected site.[73]

4.1.1.3. Decellularized matrix. Decellularization of tissue or whole organ causes the

formation of ECM protein-enriched three-dimensional bioscaffold. The acellular matrix

contains unique, tissue-specic, structural and functional molecules that regulate cellu-

lar phenotype and function, mechano-transduction, signaling, cell – matrix interactions as

well as tissue homeostasis.[74,75] Therefore, matrix decellularization followed by

recellularization with autologous cell could be a promising approach in tissue engineer-

ing. The technical dif culty of removing all cellular remnants limits the application of

decellularized matrix due to the risk of a host immune response after transplantation.Moreover, preserving the native properties of ECM, for example, three-dimensional

ultrastructure, surface topography, composition, bioactivity, and density of ligand dis-

tribution as well as internal network, such as nervous, vascular, and lymphatic net-

works, presents another technical challenge.[76] One of the important features of using

decellularized matrix is that it helps to rebuild the intricate vascular network within the

organ or tissue when the intact vascular spaces within the matrix are repopulated with

ECs.

Tissue decellularization can be accomplished by chemical, biologic, or physical

agents. Selection of an agent largely depends on tissue size, thickness, density, cellular-

ity, and lipid content. Chemical agents, for example, acids and bases, hypotonic and

hypertonic solutions, ionic, non-ionic, and zwitterionic detergents, or solvents, can be

used to decellularize tissue based on the mechanism of solubilizing cytoplasmic compo-

nents, denaturing proteins, dehydrating cells, disrupting nucleic acids, lipids, and pro-

teins. However, chemical agents may cause some disruption of the ultrastructure of

ECM as well as damage of collagen, GAGs, and growth factors.[77,78] Biologic

agents, such as nucleases, trypsin, or dispase, are basically enzymes that work by cat-

alyzing the hydrolysis of RNA and DNA chains or cleaving peptide bonds. The use of

biologic agents can also result in complications, for example, it can cause an immune

response, can disrupt ECM ultrastructure, or can remove collagen, laminin, bronectin,

elastin, and GAGs after prolonged exposure.[77,79] Physical techniques based on tem-

perature, pressure, direct force, electroporation, perfusion, agitation, pressure gradient,or super critical uid can also be used to remove cellular material from tissue. Physical



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agents work by facilitating chemical exposure, bursting cells, or disrupting cell

membranes, but can cause disruption or damage of ECM due to physical force.[77]

Repopulation of a decullarized matrix with appropriate cells is a major engineering

challenge to be solved. The parenchymal cells, for example, hepatocytes, cardiomy-

ocytes, epithelium, etc., are responsible for the specic functions of the organ, while

nonparenchymal cells, such as broblasts and ECs, help to maintain the functional

phenotype of the parenchymal cells and the cellular architecture of the tissue.[80,81]

Moreover, ECs eliminate the thrombotic barrier within decellularized matrix and protect

the parenchymal cells from the shear stress by maintaining blood ow within the

vascular network.[82,83] Therefore, the success of decellularized matrix as a scaffold

mostly depends on re-endothelialization of the intact vascular space. Autologus, allo-

genic, progenitor, or multipotent stem cells are frequently investigated for their ability

to repopulate decullarized tissues or organs. Non-immunogenic autologous cells are the

best choice because no immunosuppressive drug is required. However, harvesting com-

plexity and inadequate proliferation capability limit the use of autologous cells to

repopulate acellular matrix. By some criteria, for example, required cell population,urgency of implantation, simplicity of cell harvesting and expansion, and capability of

differentiation into desired cell types, allogenic cells could be more useful [75] if

immunogenic reaction and the use of immunosuppressive drugs can be tolerated.

Multipotent stem cells are a third option to repopulate decellularized matrix, although

challenges exist here in controlling or directing differentiation along organ-specic or

tissue-specic cell lineages.[76] Bioreactors with appropriate biophysical stimuli might

be used to assist in differentiation of stem cells.

A number of studies suggest a promising future for organ or tissue vascularization

using decellularized matrix. For example, rat hearts decellularized by coronary

perfusion and reseeded with cardiac or ECs were kept in a bioreactor, where a simu-lated cardiac physiology was maintained for 28 days. In this study, EC formed single

layers in both larger and smaller coronary vessels after day seven, and recellularized

heart exhibited pump function by day eight under physiological load with electrical

stimulation.[84] Similarly, lungs decellularized by detergent perfusion, reseeded with

epithelial or EC, and then perfused with blood and ventilated using physiologic pres-

sures resulted in gas exchange similar to that of isolated native lungs by day ve.

Moreover, when these reseeded lung constructs were reimplanted into an orthotopic

position, they generated gas exchange for up to 6 h after extubation.[85]

4.1.2. 3D printing approaches to create vascular networks

Vascular cell survival, growth, signaling, gene expression, and phenotype during matrix

tissue culture is signicantly affected by the pore architecture, pore interconnectivity

and mechanical stiffness of the scaffold.[86 – 88] The porous architecture of a scaffold

includes pore size, shape, porosity, and surface topography of the pores. Proper poros-

ity allows cell migration, proliferation, and interaction that facilitate the formation of a

vascular network. In addition, porosity assists to transport nutrients and oxygen gas to

cells and remove metabolic waste from cell surroundings. However, excessive porous

structure might result in poor mechanical strength.

In bone tissue engineering, higher porosity and larger pore size of the scaffold

enhance vascularization and tissue growth. For example, a six-day in vitro investigationon the proliferation of goat bone marrow stromal cells suggests that scaffolds having



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70% porosity and 800 μm average pore size enhance stromal cell proliferation

compared to scaffolds possessing 60% porosity and 700 μm average pore size.[89]

Like porosity, the average pore size of a scaffold has a profound effect on tissue

growth and vascularization. A number of studies suggest that variation of pore size

affects cell functionality. For example, it was reported that implanted polytetrauo-

roethylene membrane having pore size 5 μm under rat skin signicantly facilitates the

capillary blood vessel formation across the membrane – tissue interface.[90] When bro-

nectin-coated porous silicon nitride substrates seeded with 3T3 broblasts and bovine

aortic ECs were kept in ex vivo condition for ve days, EC covered the scaffold pores

having diameter below 80 μm. Interestingly, pore size under 30 μm didn’t display any

effect on EC coverage. Besides, broblast could cover pores of 50 μm and less.[91]

Further, disc-shaped porous PLA scaffolds with pore size in the range of 63 – 150 μm

facilitated vascular SMC proliferation and matrix deposition during a four-weeks cul-

ture period.[92] Therefore, a potential strategy would be to fabricate scaffolds with tar-

geted pore size that would promote the desired cellular activity, for example, tissue

ingrowth as well as blood vessel formation. The surface area of a scaffold is also animportant parameter to be considered as there exists a relationship between pore size

and available surface area for cell attachment where an RGD binding sequence motif

prevails. In general, small pore size increases specic surface area, which ensures mini-

mal ligand density for the attachment of a critical number of cells.[ 93,94] On the other

hand, large pore size facilitates cell migration but reduces cell density.

Finally, pore interconnectivity within the scaffold can have a profound inuence on

neovascularization and tissue growth. The diffusional mass transfer of nutrient and oxy-

gen gas and removal of waste metabolites can be interrupted due to improper pore

interconnectivity.[95] A β-tricalcium phosphate scaffold fabricated by the casting tech-

nique was used in a rabbit model to investigate the effect of pore parameters on vascu-larized bone tissue formation. An increase in pore size increases the size of blood

vessels, but the increase in pore interconnectivity results in an increase of size as well

as the number of blood vessels.[96] As well, pore sizes greater than 400 μm do not

support vascularization at all.[96] Scaffolds fabricated with PLGA biomaterial having a

small inter-pore distance along with minute pore size in the range of 5 – 20 μm enhances

EC growth signicantly.[97]

Physical stiffness of a scaffold can enhance formation of focal adhesions and

cytoskeletal reorganization in ECs which regulate cell migration, proliferation, and dif-

ferentiation, as well as cell – cell and cell – scaffold adhesion.[98,99] A number of studies

suggest that the spreading area of EC during focal adhesion increases with the increase

of biomaterial’s stiffness,[100,101] however, in a stiffer biomaterial, ECs like to make

cell – biomaterial interactions, rather than cell – cell interaction that cause failure to form

vascular network. On the other hand, ECs take an elongated morphology in compliant

material whose Young’s modulus is around 1 kPa, augment cell – cell interactions

instead of cell – biomaterial interaction [102], which result in the formation of

self-assembled vascular network.[100]

4.1.2.1. Rapid protyping technique and miscellaneous. To promote vascularization

within tissue engineering construct, two major factors are to be considered during scaf-

fold fabrication, one is well-interconnected microchannels or micropores, and the other

is controlled positioning of vascular cells. Some studies suggest that incorporation of channel or grooves into scaffold can promote the formation of vascular network as ECs

require aligned organization to change their function and phenotype to form



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capillaries.[103,104] More specically, the depth and topography of microgrooves has

profound effect on cell alignment and functionality. Uttayarat et al. reported that maxi-

mum alignment of bovine aortic ECs can be obtained on bronectin-coated 1 μm deep

microgrooved surface, while Hu and colleagues concluded that microgrooves having

wavy surface display better cell attachment, alignment, and survival compared to those

on the rectangularshaped microgrooved surface.[104,105]

To construct functional complex tissue with vasculature, different types of cells

should be seeded in an organized way as per their need and function, rather than in a

random fashion. Scaffolds fabricated with conventional techniques, such as gas foam-

ing, electrospinning, salt-leaching, porogen melting, molding, ber deposition, and

freeze-drying, display uneven, uncontrolled, and undesired pore size, shape, wall thick-

ness, pore interconnectivity, and morphology.[106] To fabricate the complex vascular

network layer by layer in a controlled and precise manner, CAD-based rapid prototyp-

ing (RP) techniques, such as 3D bioplotting, inkjet printing, laser-based biofabrication,

and stereolithography, can be used. However, in many cases, acellular structures were

seen to be fabricated by RP technique to keep cells free from the effect of unfavorablefabrication conditions (e.g. heat, pressure, light, adhesives, cytotoxic solvents, or mold-

ing). One major aw of acellular approach is the inef cient and inhomogeneous post-

fabrication cell incorporation. Yet, a number of studies have been accomplished to

fabricate intricate microvascular pattern with RP technique. Of several approaches, one

promising approach was to incorporate sacricial laments into scaffold by 3D bioplotter

to get complex and well-interconnected vascular pattern (Figure 6). Lewis group

demonstrated the effectiveness of fugitive ink to create 3D microvascular network by

direct write assembly method.[107] In another study, they created vascular pattern by

direct write method using fugitive ink and encapsulated the whole 3D network into

gelatin methacrylate gel (GelMA). Then HUVECs were injected after evacuation of fugitive ink. HUVECs were found to attach, proliferate, and differentiate to form vascu-

lar network.[108] To avoid cytotoxic organic solvents, and sacricial template-removal

associated extreme temperature or pressure, Wang et al. encapsulated sodium alginate-

based sacricial vascular networks fabricated with 3D bioplotter into gelatin hydrogels

and then dissolved the sacricial pattern with EDTA solution. HUVECs were injected

into evacuated channel and found to form vascular lumens and network.[109] Another

research group, Chen and colleagues fabricated a patterned 3D network composed of

sacricial carbohydrate glass lament by bioplotter, dispensed ECM (e.g. brin gels,

matrigel) encapsulated cells and mixture of poly (ethylene glycol) diacrylate and

Figure 6. 3D approaches for vascular channel formation.



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acrylate-PEG-RGDS into a rectangular mold containing the lattice lament,

photo-crosslinked the PEG hydrogel to encapsulate the 3D network as well as ECM

embedded cells, and then dissolved the lattice with cell culture media to obtain the

interconnected empty vascular lumen. HUVECs were then seeded in the vascular

lumen, and perfused with blood under high-pressure pulsatile ow. Incorporated vascu-

lar pattern was able to sustain the metabolic function of primary rat hepatocytes in the

core of a thick, densely populated tissue constructs.[110] In another study, to access the

effect of complex vascular geometry on diffusion gradient, they prepared vascular chan-

nel structure in gelatin gel by photolithography, encapsulated the microuidic network

and therapeutic cells within collagen gel, ushed gelatin gel by external ow, and then

injected HUVECs into the vascular lumen. Injected HUVECs formed conuent mono-

layers within the open microuidic channels after attachment, alignment, and prolifera-

tion. In addition, the relationship between the geometry of vasculature and spatial

diffusive gradients which is responsible for angiogenic sprouting was also demon-

strated.[111] To avoid the necrosis of large cell population in a thick tissue, vascular

pattern design parameter (e.g. vessel alignment, branching frequency and angles, andtortuosity) is required to be optimized to ensure ef cient convective diffusional mass

transfer (e.g. nutrients, metabolic wastes, gases, AFs, etc.) between cells and culture

medium or blood during in vitro or in vivo culture.[112] Therefore, several mathemati-

cal modeling and computational studies have been accomplished to optimize the archi-

tecture of vasculature considering the diffusional mass transfer, microuidic behavior,

and physiological data of human microvasculature.[113]

Controlled positioning of vascular cell into scaffold can be achieved through the

deposition of cell-laden hydrogels. Polymer hydrogels offer structural support and 3D

hydrated environment similar to in vivo condition for EC attachment, proliferation, and

differentiation as well as keep ECs free from fabrication-induced high shear forces.[114] To maintain EC viability and function in culture, hydrogel crosslinking

and scaffold fabrication process should be biocompatible. Several studies frequently

applied stereolithography,[115] inkjet printing [116,117], and 3D bioplotting [118,119]

techniques to prepare cell (e.g. ECs, SMCs) laden hydrogel to fabricate vascular pat-

tern. Besides, precise cell positioning, microchannel or interconnected porous structure

is required to avoid necrosis. To incorporate of EC-laden gel into microchannel, in a

study, liquid collagen containing EC suspension was distributed above micropatterned

poly (dimethylsiloxane) PDMS templates, the templates were then centrifuged to move

ECs into the channels, and then collagen was crosslinked to form gel. It was reported

that EC formed capillary tubes after culture with VEGF and bFGF over 24 – 48 h.[120]

While microfabrication offers many promising features to pattern microcapillary net-

work, time delay, use of nondegradable hydrogel, and handling complexity limits its

application. Therefore,an alternative way, such as mechanically removable spacer (e.g.

aligned array of wires), was investigated to create microchannel into hyrogel matrix.

Wray et al. used mechanical spacer to generate unbranched microchannel ranging from

152 to 787 μm in diameter into silk scaffolds, and then they seeded human arterial

endothelial cells (hAECs) into the hollow channels along with bioactive agents. The

formation of a contiguous layer of hAECs around the channel wall was reported

seven days after cell seeding.[121] In addition, to ensure uniform cell seeding into

incorporated microchannels, spacers can also be used to transfer self-assembled cell

layer into hydrogel matrix. In a study, casted GelMA over EC-coated micrometric goldrods positioned in a culture chamber was photo-crosslinked, an electrical potential was

applied to transfer the self-assembled layer, and then rods were removed to leave



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behind the EC-assembled layer into the matrix. The same procedure was followed to

prepare and transfer double-layered cell and gel mixture to combine microvessel stabi-

lizing 3T3 broblast cells with HUVECs.[122] Beyond non-sacricial laments,

microuidic/micromolding techniques were also investigated to fabricate sacricial pat-

tern in a microscale. In a study, sacricial gelatin meshes were prepared by micromold-

ing in patterned PDMS stamp. Hydrogels (e.g. brinogen, Type I collagen, matrigel)

were used to encapsulate the sacricial pattern, and then successive melting and ush-

ing were done to empty the interconnected channels. The viability of injected HDMECs

into microchannel was excellent as HDMECs were seen to attach, spread, and prolifer-

ate.[123] Regardless of tremendous success with sacricial templates, dissolution-

associated cytotoxic reaction limits their applications.[124] In an effort to fabricate

sacricial templates-free scaffold, Bertassoni and colleagues bioprinted agarose bers,

casted PEGDA hydrogels over the bioprinted templates, and then photo-polymerized

them. Removal of bioprinted templates was easy, as agarose didn’t adhere with casted

hydrogels. With this approach, a perfusable network could be incorporated into hydro-

gel matrix without producing any dissolution-related cytotoxicity.[125] Aside fromsacricial or removable laments, laser scanning lithography (LSL) was investigated to

create extremely precise micropattern on photo-sensitive hydrogel surface. In a study,

West group covalently incorporated RGDs and VEGF on LSL-generated micropatterned

poly (ethylene glycol) hydrogel surface. Tubules like formation in the micro-patterned

area were observed within two days, whereas no tubules formation was recognized by

day two for cells cultured on wide patterned lines.[126] As blood perfusion is very sig-

nicant to regulate vascular pattern, West and colleagues combined microfabrication

and self-assembly technique to study macro – micro-scale transport to achieve biomi-

metic perfusion in vitro. They prepared PDMS housing with uid access ports by soft

lithographic techniques, injected PEG hydrogel containing the HUVEC-10T1/2 cellmixture into the PDMS housing, photolithographically crosslinked the cell-laden hydro-

gel, and then initiated channel perfusion with normal physiologic ow rates. The perfu-

sion between the fabricated microchannel and self-assembled vascular network

signicantly facilitated the convective diffusional mass transport.[127] Moreover,

signicant success for in vivo vascularization was reported while VEGF-mimetic pep-

tide-bound PEGDA matrices were implanted into mice.[128] Besides microfabrication,

Garicia and colleagues studied functionalized hydrogel to promote vascularization

in vivo. They engineered photopolymerized PEG diacrylate hydrogel matrices contain-

ing protease-degradable sites, cell-adhesion motifs (e.g. arginine – glycine – aspartic acid),

and growth factors (e.g. VEGF) and implanted subcutaneously in male Lewis rats.

They reported that a signicant number of vessels formed into the matrices at two

weeks and the network became more densed at four weeks due to cell-demanded

sustained release of VEGF over two weeks.[129] As hydrogel density regulates the

mechanical strength, Putnam and colleagues investigated the effect of elevated brin

gel density on vasculature formation in vitro and in vivo. Their ndings suggest that

inhibition of vasculature formation in vivo due to elevated brin matrix density can be

partially lessened by co-delivery of mesenchymal stem cells (MSCs) with human

umbilical vein endothelial cells (HUVECs) into hydrogel.[130]

4.1.2.2. Microfabrication technique. Implanted, biodegradable tissue engineering scaf-

folds containing cells and growth factors successfully vascularize thin tissue (1 – 2 mm).However, vascularization of thick tissue constructs (>2 mm) is still challenging as it

takes time for blood vessels to form and anastomose with the host blood supply. To



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solve this issue, a microvascular network can be established by microfabrication within

the implantable scaffold in vitro.[131] Microfabrication and bio-microelectromechanical

techniques, are attracting attention due to their spatial resolutions of less than 10 μm,

[132] a substantial improvement over conventional scaffold fabrication techniques such

as solvent casting and porogen leaching,[133] gas foaming [134], and three-dimensional

printing.[135] However, some parameters, such as selection of fabrication material,

oxygen concentrations in the microenvironment [136], and uid shear stresses within

the microuidic scaffold [137,138], signicantly affect the success and applicability of

microfabrication techniques.

Microfabrication generally involves preparation of lithographic mask to pattern the

blood vessel network according to a model,[113] formation of a master mold by thick

photoresist or plasma etching, casting the desired biopolymer into the master mold,

incorporation of a sacricial layer to facilitate lm removal from the master mold, and

curing the lms under vacuum (Figure 7). Single-layer microuidic networks are then

stacked together to form 3D scaffolds with complex vascular microchannels before

seeding with ECs in a bioreactor.[132] Real vascular networks consist of blood vesselshaving various diameters. The direct-write laser technique can facilitate the precise

fabrication of multi-depth channels mimicking the in vivo structure.[139]

Synthetic nondegradable materials, such as poly dimethyl siloxane or silicon, are

often employed for microfabrication.[132] However, biodegradation and biocompatibil-

ity are important issues related to the integration of the vascular construct into the host

tissue as well as the potential for toxicity or inammatory responses. The biodegradable

polymers PLGA,[140] poly glycerol sebacate (PGS) [141], and silk broin [142] were

investigated for their suitability in constructing microuidic networks by microfabrica-

tion. PLGA exhibits rigid mechanical properties,[134] undesirable bulk degradation

kinetics,[143] limited biocompatibility [144], and cytotoxic degradation byproducts,[145] all of which limit the applicability of PLGA to microuidic scaffolds. PGS offers

superior mechanical properties as well as better cellular response and morphology com-

pared to PLGA.[143,144] PGS is inexpensive and attractive in terms of synthesizing

and processing simplicity to form layers up to 100 μm.[146] A third possibility,

Figure 7. Microfabrication technique for vascular network formation.



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chemically functionalized silk broin, is a promising biopolymer for microfabrication

of vascular networks [142] because it offers improved cellular adhesion, controllable

degradation, and enhanced mechanical property compared to PGS lms.

Alternative to multiple-layer stacking techniques, lithographic processes and sacri-

cial gel encapsulation methods can be applied to generate 3D microuidic networks

into hydrogels (alginate, collagen, brin, etc.) that can later be seeded with ECs for

vascular tissue engineering. A lithographic process has been used to form microuidic

channels within an alginate scaffold seeded with chondrocytes, and these channels

appear to support long-term cell survival and convective diffusion of soluble fac-

tors.[147] Examples of sacricial gel techniques include a study where Type I collagen

was used to encapsulate patterned Matrigel and dened shaped microcavities were

achieved within collagen after the digestion of matrigel by enzymes.[148] Similarly,

microuidic channels as narrow as 6 μm were formed in collagen and brin gels using

embedded sacricial gelatin meshes. The channels generated in this way facilitate the

transport of macromolecules into the channels and from channels into the gel

matrix.[123]

4.1.2.3. Modular assembly. This sophisticated tissue engineering approach achieves

formation of macro tissue constructs through the assembly of smaller modules. Micro

modules can be prepared by self-assembled aggregation,[149] microfabrication of cell-

seeded hydrogels, [150]creation of cell sheets [151], or direct printing.[152] To promote

vascularization, micro modules containing target tissue cells have been coated with ECs

and perivascular cells prior to assembly.[153] The small size of the micro modules

facilitates the transport of oxygen and metabolites even if they are seeded with high

cell density.[154] The surface of the modules can also be modied by coating with

ECM proteins. For example, bronectin coating of collagen modules implanted inimmunodecient mice increases human vascular EC survival as well as blood vessel

formation.[56]

To form macro tissue, a number of methods can be adapted to assemble the micro-

modules, such as random packing,[155] stacking of layers,[156] or directed assem-

bly.[157] In the random packing approach, EC-coated modules are packed together into

a larger chamber and perfused with blood or culture medium which facilitates the

formation of interconnected channels among the interstitial spaces between modules

(Figure 8). Antithrombogenic ECs assemble to form vascularized tissue, and promote

functional perfusion by delaying clotting time and inhibiting loss of platelets.[153]

Incorporation of stem and progenitor cells along with ECs can signicantly affect the

stabilization of blood vessels that develop when using the modular approach. For exam-

ple, implanted collagen gel modules coated with rat aortic ECs into Sprague-Dawley

rats caused the formation of blood vessels within the rst seven days. The nascent

blood vessels eventually became mature and then lasted at least for 60 days. In these

simple constructs, the new blood vessels connected with the host vasculature, yet some

of them were leaky.[158] Interestingly, addition of bone marrow-derived stem cells to

the EC coated modules lead to formation of blood vessels having similar density, but

less leakiness.[159] However, the random-packed modular approach causes lack of

mechanical integrity in the resultant macro tissue and limits the use of secondary cells

other than ECs.

Sequential assembly is a more directed method used to construct large tissue frommodular tissues with a specic microarchitecture.[160] In this approach, microgels con-

taining specic architectural designs are assembled in a controlled fashion to connect



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the micro channels of each microgel, resulting in the formation of a bifurcating and

interconnected network.[112] Thus, sequential assembly provides better control over

the relative spatial arrangement of the building blocks. For example, concentric PEG

microgel building blocks embedded with network of microchannels were assembled

sequentially by photo-crosslinking to achieve a natural blood vessel-like structure. The

inner layer was seeded with HUVECs, and the outer layer was coated with SMCs to

achieve a tube-like structure using a mineral oil immersion technique.[1] Formation of

vascular network by 3D fabrication technique has been summarized in Table 3.

4.1.2.4. Comparison among different fabrication techniques. It is well established that

oriented pores or microchannels are required to maintain cell viability in tissue engi-

neering construct. In previous section, it was already stated that 3D fabrication tech-

niques are associated with unfavorable conditions which led researchers to fabricate

acellular construct. However, post-fabrication cell seeding into scaffold demonstrates

many complexities in terms of ef cient and homogeneous cell positioning. In contrast,

fabrication of cellular scaffold facilitates the controlled and homogeneous cell distribu-

tion. Compared to other fabrication material, hydrogels are more suitable to fabricate

cellular scaffold as it can ensure cell viability by providing in vivo like environment.

However, lack of mechanical strength limits the applications of hydrogel scaffold. In

addition, free radical photo polymerization or chemical crosslinking to polymerize

hydrogels can negatively affect the viability of encapsulated cells. Channel, which isthe key parameter to promote mass transport and directional growth of EC, can be

introduced into the 3D scaffold by using sacricial and non-sacricial laments.

Figure 8. Modular assembly technique for vascular network formation.



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T a b l e 3 .

S c a f f o l d f a b r i c a t i o n t e c h n i q u e s f o r v a s c u l a r i s a t i o n .

F a b r i c a t i o n

t e c h n i q u e

C h a n n e l / p o r e / p a t t e r n

/

s p a c e r

G e l a t i o n

m e t h o d

S c a f f o l d m a t e r i a l

A d d i t i v e s / a d d i t i o n

t e c h n i q u e

M o d e l / t a r g e t

s i t e

C h a n n e l c o n u e n c e / l u m e n

f o r m a t i o n

R e f

3 D B i o p l o t t e r

S a c r i c i a l

l a m e n t s / p l u r o n i c

F 1 2 7

P h o t o

G e l M A

H U V E C s

I n v i t r o

F o r m e d c o n u e n t l a y e r , v

i a b i l i t y

> 9 5 % a f t e r 4 8 h

[ 1 0 8 ]

B u l k p o l y m e r i z a t i o n

M e c h a n i c a l s p a c e r

P h o t o

G e l M A

H U V E C s

I n v i t r o

C o n u e n t l u m e n o f 1 0 m m l o n g

a n d 6 1 8 ± 1 5 μ m d i a m e t e r ,

s t a b l e g e o m e t r y f o r 1 5 d a y s

[ 1 2 2 ]

3 D B i o p l o t t e r

A g a r o s e s p a c e r

T h e r m a l

M u l t i c e l l u l a r

s p h e r o i d s ,

d i a m e t e r 3 0 0 –

5 0 0 μ m

C H O ,

H U V S M C s ,

H S F s

I n v i t r o

S p h e r o i d s f u s e d w i t h i n 5 – 7 d a y s

t o f o r m v a s c u l a r t u b e s o f

0 . 9 –

2 . 5 m m d i a m e t e r

[ 3 2 8 ]


S a c r i c i a l

l a m e n t s / c a r b o h y d r a

t e

g l a s s

I o n i c ,

p h o t o ,

t h e r m a l ,

e n z y m a t i c

A l g i n a t e , b r i n ,

p e g d a , m a t r i g e l

H U V E C s

I n v i t r o

C o n u e n t l a y e r w i t h i n o n e d a y

[ 1 1 0 ]


S a c r i c i a l l a t t i c e

/ s o d i u m a l g i n a t e

T h e r m a l ,

i o n i c

G e l a t i n , a g a r o

s e

a n d c o l l a g e n

H U V E C s

I n v i t r o

C o n u e n t l a y e r o n g e l a t i n

a n d

c o l l a g e n , s p h e r o i d s o n a g a r o s e

a f t e r 3 – 4 d a y s

[ 1 0 9 ]

3 D M i c r o m o l d i n g

N o n s a c r i c i a l a g a r o

s e

l a m e n t s

P h o t o ,

t h e r m a l

1 0 % G e l M A

H U V E C s

I n v i t r o

C o n u e n t l a y e r b y d a y s e

v e n

[ 1 2 5 ]

S t e r e o l i t h

o g r a p h y

H e x a g o n a l / w o o d p i l e

m i c r o p a t t e r n

P h o t o

1 0 o r 1 5 %

G e l M A

H U V E C s /

i m m e r s i o n a n d

a g i t a t i o n

I n v i t r o

C o n u e n t l a y e r b y d a y f o

u r

[ 3 2 9 ]

3 D R o b o

t i c

d i s p e n s i n g

5 % g e l a t i n l a m e n t ,

c h a n n e l 4 0 0 – 7 0 0 μ m

C h e m i c a l

3 – 9 m g / m L

c o l l a g e n

H U V E C s / m i x i n g

I n v i t r o

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