additive manufacturing techniques for the production of...

Additive manufacturing techniques for theproduction of tissue engineering constructsCarlos Mota, Dario Puppi, Federica Chiellini and Emo Chiellini*Laboratory of Bioactive Polymeric Materials for Biomedical and Environmental Applications (BIOLab), UdR-INSTM, Department ofChemistry and Industrial Chemistry, University of Pisa, San Piero a Grado (Pi), Italy

Abstract

‘Additive manufacturing’ (AM) refers to a class of manufacturing processes based on the building of asolid object from three-dimensional (3D) model data by joining materials, usually layer upon layer.Among the vast array of techniques developed for the production of tissue-engineering (TE)scaffolds, AM techniques are gaining great interest for their suitability in achieving complex shapes andmicrostructures with a high degree of automation, good accuracy and reproducibility. In addition, thepossibility of rapidly producing tissue-engineered constructs meeting patient’s specific requirements, interms of tissue defect size and geometry as well as autologous biological features,makes them a powerfulway of enhancing clinical routine procedures. This paper gives an extensive overview of different AMtechniques classes (i.e. stereolithography, selective laser sintering, 3D printing, melt–extrusion-basedtechniques, solution/slurry extrusion-based techniques, and tissue and organ printing) employed forthe development of tissue-engineered constructs made of different materials (i.e. polymeric, ceramicand composite, alone or in combination with bioactive agents), by highlighting their principles and tech-nological solutions. Copyright © 2012 John Wiley & Sons, Ltd.

Received 29 June 2012; Revised 2 August 2012; Accepted 27 September 2012

Keywords additive manufacturing; solid freeform fabrication; tissue engineering; regenerative medicine;tissue and organ printing; scaffold

1. Introduction

The emergence in the last two decades of novel regenera-tive approaches in the biomedical field has raised the needfor new technologies employable in the processing ofbiodegradable polymeric materials, with tailored physico-chemical structural features and degradation properties.This has led to the development of a range of novel tech-niques and methodologies enabling the fabrication ofmicro- and nano-structured biodegradable constructs suit-able for applications in tissue engineering (TE) and regener-ative medicine.

TE typically involves the use of what is commonlyreferred to as a ‘scaffold’ that serves as a temporary template

for interactive trafficking of cells and for the formation ofthe extracellular matrix (ECM), thus providing structuralsupport for the newly formed tissue (Mano et al., 2007). Asuccessful scaffold should meet some basic requirements,generally involving biocompatibility, biodegradability withcontrolled kinetics, interconnected porous structure with atailored pore size, mechanical properties close to the targettissue and predefined geometry and size.

A broad variety of biomaterials have been tested for TEscaffolds, ranging from natural polymers showing biologi-cal cues suited to promote desirable cell responses, to syn-thetic polymers with more controllable physicochemical,mechanical and processing properties (Place et al., 2009;Puppi et al., 2010a). Recent evidence suggests that, besidesmaterial chemistry, scaffold micro- and nanostructural fea-tures affect cell adhesion, spreading, growth, propagationand reorganization (Leong et al., 2003; Karageorgiou andKaplan, 2005; Stevens and George, 2005; Puppi et al.,2010b; Mota et al., 2011). Depending on scaffolding mate-rial and TE strategy, different processing techniques andmethodologies have been proposed to optimize final

*Correspondence to: E. Chiellini, Laboratory of BioactivePolymeric Materials for Biomedical and EnvironmentalApplications (BIOLab), UdR-INSTM, Department of Chemistryand Industrial Chemistry, University of Pisa, Via VecchiaLivornese 1291, 56010 San Piero a Grado (Pi), Italy. E-mail:[email protected]

Copyright © 2012 John Wiley & Sons, Ltd.

JOURNAL OF TISSUE ENGINEERING AND REGENERATIVE MEDICINE REVIEWJ Tissue Eng Regen Med 2012.Published online in Wiley Online Library (wileyonlinelibrary.com) DOI: 10.1002/term.1635

scaffold performances in terms of external shape andsize, surface morphology and internal architecture. Theseinclude, among others, solvent casting combined withparticulate leaching, freeze drying, gas foaming, meltmoulding, fibre bonding, phase separation techniques,electrospinning and additive manufacturing (AM) tech-niques (also known as solid freeform fabrication techniques)(Puppi et al., 2010a).

AM is defined as ‘the process of joining materials tomake objects from three-dimensional (3D) model data,usually layer upon layer’ (ASTM, 2012), and it has beenextensively applied for the fabrication of TE scaffolds bymeans of different techniques, such as stereolithography(SLA) and fused deposition modelling (FDM) (Woodruffand Hutmacher, 2010). These techniques enable 3Dstructures with a predefined geometry and size to beobtained, and with a porous architecture characterized bya fully interconnected network of pores with customizablesize, shape and distribution. 3D model data for scaffolddevelopment can be derived from medical imaging tech-niques used for diagnostic purposes, such as computertomography (CT) and magnetic resonance imaging (MRI),and are generally treated by computer-aided design(CAD) and computer-aidedmanufacturing (CAM) software(Sun and Lal, 2002; Hieu et al., 2005) (Figure 1). Alterna-tively, a simplified 3D model can be directly designed inCAD software or developed by means of mathematical

equations (Gabbrielli et al., 2008) or topological optimiza-tion (Almeida and Bártolo, 2010). Most CAD softwareconverts the 3Dmodel into a standard tessellation language(STL) file containing the information of the 3D object’ssurface geometry. The STL file is then sliced into layersoriginating a slice file (SLI) that is loaded digitally into themachine that drives the motions of the build parts.

Depending on the fabrication principle, the most exten-sively applied AM techniques for TE scaffold fabricationhave been conventionally classified into four categories:(a) stereolithography (SLA); (b) selective laser sintering(SLS); (c) three-dimensional printing (3DP); and (d)fused deposition modelling (FDM) (Hutmacher et al.,2004). However, a number of AM techniques based eitheron well-known industrial technologies or on innovativeprinciples are being introduced in the TE field andadapted to meet the requirement of materials processingfor scaffold fabrication. The wide range of innovativeAM techniques recently developed has often led todebatable terminology for the definition of techniquescombining principles inherent to different classes of theaforementioned classification (Figure 2).

This review is dealing specifically with AM techniques,currently investigated for the development of polymeric,ceramic and composite scaffolds or cell-encapsulatingliving constructs for TE applications, by highlighting theirbasic principles and open issues (Table 1).

Figure 1. Schematic representation of some basic steps in the scaffold-based TE approach, involving: (a) data acquisition by themedical imaging technique; (b, c) 3D computer solid model of tissue defect and biomimetic scaffold; (d, e) layer-by-layer 3D scaffoldmanufacturing; (f, g) cell seeding and dynamic cell culture of tissue-engineered construct; (h, i) scaffold implantation and tissueregeneration

C. Mota et al.

Copyright © 2012 John Wiley & Sons, Ltd. J Tissue Eng Regen Med 2012.DOI: 10.1002/term

2. Stereolithography (SLA)SLA is an AM technique that uses an ultraviolet (UV) lightor laser to selectively polymerize layers of a photosensitivepolymer (Figure 3a). The first studies on the production of3D objects by selective photopolymerization of a liquidresin using UV light were performed at the beginning ofthe 1980s by Kodama (1981), who developed twoapproaches, one using masks for the definition of eachlayer structure, the other involving the use of an opticalfibre to selectively polymerize the liquid resin. In the lattercase, a predefined pattern was achieved by controlling themovement of the fibre in the X and Y axes. In 1986, Hullattributed the denomination of SLA to the process ofproducing 3D solid objects layer-by-layer by means of UVlight (Hull, 1986, 1990).

After the polymerization of a layer, the constructionplatform is lowered a given distance and a recoat bar placesa new uniform layer of resin on top of the previously builtone. In order to prevent delamination between adjacentlayers, the polymerization of a layer is carried out byoverlapping a percentage of the previous built layer. Thelayer-by-layer process is repeated until the 3D object isfinally built up. Further processing steps include theremoval of the non-polymerized resin and post-curing ofthe green part in order to improve polymerization betweenlayers and to reduce surface irregularities (Hutmacheret al., 2004; Melchels et al., 2010b).

A recently developed SLA technique involves a digitallight projector source to direct light using a DigitalMicromirror Device™ (DMT™) constituted by an array ofmirrors that selectively divert the light to the vat containingthe photopolymerizable polymer (Melchels et al., 2010b).This technique involves the use of a smaller quantity of pho-topolymer when compared to conventional SLA equipment.

The preparation of photocrosslinkable liquid resinsgenerally requires toxic solvents (Elomaa et al., 2011),and the commercially available resins (epoxy-based oracrylate-based) suitable for SLA processing present lackof biocompatibility and biodegradability (Jansen et al.,

2008; Melchels et al., 2010b). For such reasons, the firstattempts to use SLA technology for the production ofbiomedical implants were made by carrying out anindirect approach. Levy et al. (1999) employed a slurrymade from a suspension of hydroxyapatite (HA) powderin a liquid photocurable acrylic resin. After the lasercuring process, the acrylic resin was removed by heatingin order to obtain a porous HA prosthesis. A ‘lost-mould’approach was developed by Chu et al. (2001), whoinfiltrated an epoxy mould with a HA suspension inacrylate binders; afterwards the mould and the binderswere removed by pyrolysis, followed by a sintering process.The direct use of SLA technology to produce scaffolds wasonly possible after the development of new biocompatibleand biodegradable materials. However, the amount ofphotopolymerizable biomaterials processed using SLAwithproperties suitable for TE applications is still limited(Hutmacher et al., 2004). Recently, new photocurablepolymers with enhanced properties for scaffold productionhave been developed. Polymer networks with tunablehydrophilicity and mechanical properties, prepared byphotocrosslinking fumaric acidmonoethyl ester functionalizedwith three-armed poly(D,L-lactide) (PDLLA) oligomers andusing N-vinyl-2-pyrrolidone as diluent and co-monomer,showed good cytocompatibility during in vitro cell cultureexperiments employing mouse preosteoblasts (Jansenet al., 2008). Using SLA, porous structures with well-defined gyroid architecture were prepared from these novelphotocurable polymers. A recent study reported on thepreparation of porous polylactide constructs by SLAwithoutthe use of reactive diluents (Figure 4a, b) (Melchels et al.,2009). Star-shaped PDLLA oligomers were functionalizedusing methacryloyl chloride and photocrosslinked in thepresence of ethyl lactate as a non-reactive diluent, enablingthe production of porous scaffolds and films (Melchels et al.,2009). The films were seeded with murine preosteoblaststhat adhered and proliferated well, although cell viabilitywas significantly lower than controls grown on tissue culturepolystyrene. The previously mentioned PDLLA-based resinwas used for the production of scaffolds with gyroid

Figure 2. Proposed classification of the AM techniques commonly employed in TE

Additive manufacturing techniques for the production of tissue engineering constructs


Table

1.Prosan

dco

nsofAM

tech

niques

curren

tlyem

ploye

dforth

eproductionoftissue-en

ginee

ringco

nstructs

Tech

niqu

eMaterials

Resolution

(mm)

Adv

antage

sDisad

vantag

esRe

ferenc

es

Powde

r-ba

sedtech

niqu

es

SLS

Polymer,c

eram

icor

compo

sites

50–10

00Nosupp

ortmaterialn

eede

dLaserintensitycanindu

cepo

lymer

degrad

ation;

scaffoldsge

nerally

presen

tlow

mecha

nicalp

rope

rties,

limited

andhigh

-costmaterials,

trap

pedno

n-sintered

material,po

orco

ntrolo

versurfacetopo

grap

hy

(Tan

etal.,20

03;C

iardelliet

al.,20

05;

Williamset

al.,20

05;W

iria

etal.,20

07;

Dua

nan

dWan

g,20

10a,

2010

b;Dua

net

al.,

2010

;Eosolyet

al.,20

10;Y

eong

etal.,20

10;

Zhou

etal.,20

10;D

uanet

al.,20

11a,

2011

b;Krishn

aet

al.,20

11;S

huai

etal.,20

11)

3DP

Polymer,c

eram

icor

compo

sites

50–30

0Nomaterialshe

atinginvo

lved

inthefabricationproc

ess,no

supp

ortmaterialn

eede

d,high

prod

uction

rate,p

ossibilityof

prod

ucinglargesize

samples

Structureshrink

ingup

onsintering,

trap

pedno

n-bo

undmaterial,low

scaffold

mecha

nicalp

rope

rties,high

roug

hnessof

scaffold

surface

(Leu

kers

etal.,20

05;S

eitz

etal.,20

05;

Gbu

reck

etal.,20

07;W

illet

al.,20

08;

Shan

jani

etal.,20

10;B

utsche

ret

al.,20

11;

Meszaroset

al.,20

11;T

arafde

ret

al.,20

12)

Photosen

sitiv

ity-ba

sedtech

niqu

es

SLA

Polymers,hy

drog

els

14–15

0Re

latively

easy

toremov

esupp

orts,p

ossibilityof

encaps

ulatingcells

Necessity

ofremov

ingno

n-po

lymerized

resin,

post-curingge

nerally

requ

ired

,shrink

ageof

thescaffoldsdu

eto

post-

proc

essing

step

s,few

suitab

lebioc

ompa

tiblean

dbiod

egrada

ble

materials

(Dha

riwalaet

al.,20

04;A

rcau

teet

al.,20

06;

Jansen

etal.,20

08;M

elch

elset

al.,20

09,

2010

a;Elom

aaet

al.,20

11)

mSLA

Polymers,

compo

sites

0.5–

10Highaccu

racy,smalla

mou

ntof

photosen

sitive

material

necessary

Limited

overallscaffoldsize

(Luet

al.,20

06;C

hoie

tal.,20

09;L

anet

al.,

2009

;Lee

etal.,20

09)

Two-ph

oton

mSLA

Polymers

0.1–

4Po

ssibility

ofprod

ucing

scaffolds

witho

utlayers

supe

rimpo

sing

,high

accu

racy

Limited

overallscaffoldsize,d

istortion

andshrink

agecanoc

cur

(Ovsianiko

vet

al.,20

07;C

laeyssen

set

al.,

2009

;WeiÃet

al.,20

09)

Melt–extrusion-ba

sedtech

niqu

es

(FDM,P

ED,B

ioplotter,

Bioe

xtrude

ran

dBioscaffolde

r)

Polymer,c

ompo

sites

100–

500

Goo

dmecha

nicalp

rope

rties,no

materialstrap

pedinto

the

scaffold

Possible

thermal

degrad

ation

ofpo

lymers,

relative

regu

larstructure,

supp

orts

need

edforstructureov

erha

ngs

(Hutmache

ret

al.,20

01;M

oron

ietal.,20

06;

Gloriaet

al.,20

09;R

agae

rtet

al.,20

10;

Dom

ingo

set

al.,20

11;H

amid

etal.,20

11;

Sobral

etal.,20

11;Y

ilgor

etal.,20

12)

Solution

/slurryextrusion-ba

sedtech

niqu

es

PAM/PAM2

Polymers(PAM),

hydrog

elsan

dcells

(PAM2)

10–10

00Po

ssibility

ofprod

ucing

hydrog

elscaffoldsen

caps

ulatingcells

(PAM2)

Organ

icsolven

ts(PAM)

(Voz

ziet

al.,20

02,2

003;

Tirella

etal.,20

11b)

LDM

Com

posite

slurries

300–

500

Highaccu

racy

onde

position

ofpo

lymericslurries,

micropo

rosity

dueto

free

zedrying

Organ

icsolven

ts(Xiong

etal.,20

02;X

uet

al.,20

10)

Robo

casting

Ceram

icsor

compo

sites

100–

450

Highlyco

ncen

trated

collo

idal

susp

ension

scanbe

used

Sinterizationne

cessaryin

mostcases

(Miran

daet

al.,20

06;R

ussias

etal.,20

07;

Miran

daet

al.,20

08;F

ranc

oet

al.,20

10;

Martine

z-Vazqu

ezet

al.,20

10)

C. Mota et al.


architecture that were characterized in comparison with arandom pore architecture resulting from salt leaching(Melchels et al., 2010a). After dynamic seeding followedby 5days of static culture, gyroid scaffolds showed large cellpopulations in the inner part of the scaffold, while salt-leached scaffolds were covered with a cell sheet on theoutside and no cells were found in the scaffold centre.

UV-crosslinked natural or synthetic hydrogels are oftenemployed in combination with cells for the engineering ofsoft tissues, because of their ability to trigger cellresponses and induce the formation of new tissue (Tsangand Bhatia, 2004; Fedorovich et al., 2007). Severalfactors, such as UV light intensity, exposure time and freeradicals that can be generated from the photoinitiator,determine the success of cell encapsulation (Lu et al., 2006).As an example, poly(ethylene oxide) and poly(ethyleneglycol)dimethacrylate photopolymerizable hydrogels wereused for the production of constructs encapsulating chinesehamster ovary cells by means of SLA, achieving good cellvitality by minimizing the amount of the photoinitiatior thatcould have toxic effects (Dhariwala et al., 2004). Furtherimprovements carried out by Arcaute et al. (2006) on poly(ethylene glycol) dimethacrylate hydrogels allowed anincrease in the survival of encapsulated human dermalfibroblasts observed after 24h.

Considerable drawbacks of SLA are represented by theshrinkage of the scaffold structure during the productionprocess and the post-processing step, the necessity ofauxiliary supports for the construction of complexgeometries, and the difficulty of loading bioactive agentsand obtaining multi-material structures (although newapproaches involving the combination of vats containingdifferent resins have recently been proposed (Wanget al., 1996; Choi et al., 2011). In addition, SLA has atypical resolution in the range of 150 mm in the threedirections of space. While the vertical resolution (alongthe built axis) is mainly related to the chemistry of theprocess, and therefore can be enhanced by modifyingthe photosensitive resins, major improvements in thelateral resolution (in-plane) are related to the resolutionof the scanning system used for resin surface irradiationand to the quality and stability of the light beam (Bertschand Renaud, 2011). A number of microstereolithography(mSLA) techniques have evolved from SLA to improvethe resolution of the manufacturing process. Theseinclude constrained-surface and free-surface mSLAtechniques that use a vector-by-vector tracing of eachlayer of the object, and integral mSLA techniques thatare based on the projection of a high resolution imagecreated by a dynamic mask on the surface of the photo-sensitive resin. For instance, poly(propylene fumarate),a biodegradable and UV photocurable material, wasinvestigated for the production of scaffolds for bone TEby means of different mSLA techniques (Choi et al.,2009; Lan et al., 2009; Lee et al., 2009). In addition,sub-micron mSLA processes involving the manufacturingof the object directly inside the reactive medium, withoutlayers superimposing, have recently been developed.Among them, two-photon mSLA allows size resolutionsTa

ble

1.(Continued

)

Tech

niqu

eMaterials

Resolution

(mm)

Adv

antage

sDisad

vantag

esRe

ferenc

es

Tissue

andorga

nprinting

Mecha

nical

depo

sition

/disp

ensing

Hyd

roge

lsan

dcells

600–

1000

Largesize

base

construc

tscan

beprod

uced

Limited

accu

racy

onstrand

depo

sition

,lim

ited

thickn

essan

dpo

rosity

due

tofibres

colla

psing

(Fed

orov

ichet

al.,20

11a,

2011

b)

Ink-jettech

nologies

Cells,p

roteinsan

dhy

drog

els

60–10

0Highsurvival

ratesafterprinting

proc

ess

Cella

gglomerationor

sedimen

tation

inside

thecartridg

e(Bolan

det

al.,20

06,2

007;

Saun

ders

etal.,

2008

;Cui

andBo

land

,200

9;Tirella

etal.,

2011

a)LaBP

/BioLP

™Cellsan

dhy

drog

els,

ceramics

50–50

0Dep

ositionof

high

cellde

nsities

andsm

allv

olum

eswithhigh

accu

racy;a

widerang

eof

cell

lines

werealread

yprinted

achievinghigh

cellvitality

Risk

ofdrop

lets

splashing

(Barronet

al.,20

04;O

thon

etal.,20

08;

Guille

mot

etal.,20

10;G

uillo

tinet

al.,20

10;

Catroset

al.,20

11;G

ruen

eet

al.,20

11;K

och

etal.,20

12;P

irlo

etal.,20

12)



down to 100 nm to be achieved by employing a femtosec-ond laser (WeiÃ et al., 2009). The first study proposingtwo-photon mSLA for the development of polymericporous structures was carried out by Ovsianikov et al.(2007), who investigated two commercially availablematerials, ORMOCERW (hybrid organic–inorganic poly-meric material) and SU8 (epoxy-based polymer). Thedeveloped 3D structures showed good in vitro cytocom-patibility, but their application for TE purposes was

limited by the lack of biodegradability. New biodegrad-able materials suitable for processing by means oftwo-photon mSLA have recently been developed. Forinstance, the biodegradable triblock co-polymer poly(e-caprolactone-co-trimethylenecarbonate)-b-poly(ethyleneglycol)-b-poly(e-caprolactone-co-trimethylenecarbonate)combined with 4,40-bis(diethylamino)benzophenoneas photoinitiator was recently employed to producescaffolds with a resolution of 4 mm, although the

Figure 3. Schematic representation of the most commonly AM techniques employed in TE. (a) Stereolithography (SLA): layers ofa photosensitive polymer are selectively polymerized by a laser or UV light; (b) selective laser sintering (SLS): a powder bed isselectively sintered using a high intensity laser beam, and after the generation of a layer, a new powder bed is spread mechanicallyby a roller; (c) 3D printing (3DP): using an inkjet head, a binder material is selectively laid on a powder layer, then a new layer ofpowder is placed on top of the previous one and the process of deposition restarts; (d) fused deposition modelling (FDM): apolymeric filament is extruded through a heated nozzle, and the polymer melt is continuously deposited on a construction platformto build the scaffold layer-by-layer; (e) solution/slurry extrusion-based techniques: a solution or slurry is continuously extruded witha predefined pattern onto a construction platform; (f) tissue and organ printing techniques: cells and bioactive agents are deposited,usually in the form of self-assembling tissue building blocks, obtained by employing hydrogels to form biopaper or bioink (example ofmechanical dispensing of cell-containing bioink onto a hydrogel biopaper)

C. Mota et al.


developed structure suffered from distortion, mostlikely due to polymer shrinkage during polymerization(Claeyssens et al., 2009).

3. Selective laser sintering (SLS)Selective laser sintering (SLS) was developed and patentedby Deckard (1989) and is based on the selective sintering of

a polymer, ceramic or hybrid powder bed, using a high-intensity laser beam (e.g. CO2 laser) (Figure 3b). After thegeneration of a layer by selective particle bonding, a newpowder bed is spread mechanically by a roller on top ofthe previous one to build up the 3D object layer-by-layer.The sintering process is performed with proper laserintensity capable of bonding the particles in a layer, andalso between adjacent layers, in order to form a cohesive3D structure. The dissociated or non-sintered areas serveas supports for the subsequent layers.

The mechanical properties and size accuracy of thescaffolds prepared by SLS are strongly dependent on theconstruction plane and processing parameters, such asmanufacturing direction, scan spacing, size of the particlesand laser intensity (Ciardelli et al., 2005; Williams et al.,2005; Eosoly et al., 2010). Generally, the scaffolds producedwith this technique present lowmechanical properties, thuslimiting their application to non-loadbearing sites (Eosolyet al., 2010; Duan et al., 2011a; Krishna et al., 2011).However, studies using the SLS technique generally aimfor the production of scaffolds for bone TE, althoughrecently the production of scaffolds for cardiac tissueregeneration was also investigated (Yeong et al., 2010).

Initial experiments on the production of TE scaffolds bySLS were performed in 2003, employing a blending ofpolyetheretherketone and HA (Tan et al., 2003). Subse-quently, anatomically shaped poly(e-caprolactone) (PCL)scaffolds were successfully fabricated by SLS, showingmechanical properties within the lower range of trabecularbone and the ability to support in vivo bone tissue growth(Williams et al., 2005). Further studies reported thedevelopment of PCL–HA composite scaffolds by SLS, dem-onstrating that it is possible to control scaffold morphologyby changing the processing parameters (Wiria et al., 2007;Eosoly et al., 2010).

Poly(L-lactide) (PLLA)-based composite scaffolds werealso produced by processing a bed of PLLA or PLLA–carbonated HA nanocomposite microspheres (Zhou et al.,2010). However, the developed scaffolds presented anirregular morphology, due to non-homogeneous fusionof nanocomposite microspheres, but they were able tosupport the in vitro adhesion of human osteoblast-likecells.

A number of recent studies have reported on thedevelopment of poly(hydroxybutyrate-co-hydroxyvalerate)–tricalcium phosphate (PHBV–TCP) scaffolds by applyingSLS to composite microspheres (Figure 4c, d) (Duan andWang, 2010a, 2010b; Duan et al., 2010, 2011a, 2011b). Thebioactivity of the composite scaffolds was enhanced byloading the microspheres with bovine serum albumin(BSA) as model protein (Duan and Wang, 2010b) or byphysically entrapping gelatine on the scaffold surface (Duanet al., 2011b) and binding human bone morphogeneticprotein-2 (BMP-2) to heparin immobilized on gelatine-modified scaffold surface (Duan and Wang, 2010a).

Recent studies have proposed SLS for the manufacturingof bioactive ceramic scaffolds. For instance, bioactiveglasses were processed in combination with stearic acidused as binder (Krishna et al., 2011). After removal of

Figure 4. Scaffolds produced by means of different AM techniques:(a) photograph and (b) scanning electron microscopy (SEM) imageof polylactide scaffold by stereolithography (SLA) (Melchels et al.,2009); (c) photograph and (d) SEM image of poly(hydroxybutyrate-co-hydroxyvalerate)–tricalcium phosphate (TCP) nanocompositescaffold by selective laser sintering (SLS) (Duan et al., 2011a);(e) photograph and (f) SEM image of TCP scaffold by 3DP (Tarafderet al., 2012); (g) photograph and (h) SEM image of poly(«-caprolactone) (PCL) scaffold by the melt–extrusion-basedtechnique (Bioextruder) (Domingos et al., 2009); (i) 3D mCTmodel and (j) SEM image of bioactive glass scaffold by the solutionextrusion-based technique (direct ink-writing) (Fu et al., 2011)



stearic acid, the scaffolds presented a pore size in the range300–800mm and an average porosity of 50%. In vitro cellculture experiments with mouse osteoblasts showed anincrease during 6days of cell proliferation over the scaffoldsection. Recently, a binder-free approach involving thedirect sintering of HA nanoparticles was shown to besuitable for the fabrication of simple-structure scaffolds(Shuai et al., 2011).

Hybrid scaffolds combining synthetic and naturalpolymers by SLS are not commonly investigated, due tothe fast degradation of the natural components whensubmitted to laser beam treatment. However, scaffoldscomposed of PCL and different polysaccharides (starch,gellan and dextran) showed only superficial degradationof the polysaccharides, while the bulk remained unchanged(Ciardelli et al., 2005). The incorporation of gellan andstarch allowed cell adhesion on the produced constructsto be enhanced.

Despite of the advantages of SLS (e.g. no need ofconstruction support), biomaterials in the form of finepowder suitable to be processed by this technique arelimited and their cost is high (Zhou et al., 2010).Moreover, most methodologies for the preparation ofpolymeric and ceramic particles are still performed atlaboratory scale for non-commercial purposes. Otherconcerns associated with this technique are the difficultyof eliminating the entrapped material inside complexgeometries, and the poor control over surface topographythat is generally defined by particle size and geometry(Ciardelli et al., 2005; Butscher et al., 2011).

4. Three-dimensional printing (3DP)

Three-dimensional printing (3DP), originally developedat the Massachusetts Institute of Technology at thebeginning of the 1990s (Sachs et al., 1990, 1993), is basedon the controlled deposition of a binder material laid on apowder layer using an inkjet head (Figure 3c). After theselective deposition of the binder, a new layer of powderis placed on top of the previous one and the process ofdeposition restarts. The unbound powder that composeseach layer acts as support for the object being built. Afterthe fabrication of each layer, the construction platformcontaining the powder layer moves downward a distanceequivalent to the thickness of the layer deposited. Thelayer-by-layer process goes on until the 3D object is finallybuilt. The high production rate and the possibility ofobtaining large-sized models at a low cost make thisprocess interesting for industrial applications.

Polymeric, ceramic and composite powder materialshave been investigated for the production of TE scaffoldsby 3DP (Butscher et al., 2011). The first published paperon TE scaffolds by 3DP was reported by Kim et al.(1998), who used a mixture of poly(lactic-co-glycolicacid) (PLGA) and NaCl as the powder bed. By employinga solvent to bind the polymeric particles, uniformlydistributed pore channels with a size of 800 mm were

obtained. After salt leaching, the structural elementspresented pores sized in the range 45–150 mm. Duringco-culture experiments employing hepatocytes and non-parenchymal cells, good cell attachment and albuminsynthesis in both static and dynamic conditions wereobserved.

Natural polymers have been also proposed for scaffoldfabrication by 3DP. For instance, starch was combinedwith dextran and gelatin to produce scaffolds, using wateras binder (Lam et al., 2002). However, these starch-basedscaffolds presented limited structural integrity and lowmechanical properties after drying at 100�C. Slightimprovements of the mechanical properties wereachieved by infiltration of the scaffolds with a PLLA–PCLpolymeric solution.

Bioactive ceramics, such as HA, bioactive glasses andtricalcium phosphates (TCP), are the most commonmaterials used for scaffold production using the 3DP tech-nique (Figure 4e, f) (Leukers et al., 2005; Seitz et al.,2005; Meszaros et al., 2011; Tarafder et al., 2012). b-TCPbone fillers, produced by a patented 3DP technique knownas Theriform, were among the first medical productsfabricated by an AM technique (Donald et al., 2002). Thecombination of different ceramic materials, such as TCPand tetracalcium phosphate (TTCP), was investigated forthe development of drug release matrices (Gbureck et al.,2007). Generally, post-processing steps (e.g. sintering)are needed to improve implant mechanical properties(Butscher et al., 2011). In order to enhance the binding ofceramic particles, a 3DP technique involving the processingof a mixture of polymer and ceramic particles and theelimination of the polymer phase after sintering wasinvestigated (Shanjani et al., 2010). Several types ofpolymeric powders acting as binder were investigated. Forinstance, starch was blended with HA particles andprocessed employing an aqueous binder (Will et al., 2008).A further study proposed the processing of poly(vinylalcohol)/tricalcium polyphosphate (PVA/TCP) blends incombination with an aqueous binder (Shanjani et al., 2010).

The major disadvantage of the 3DP technique isshrinkage of the scaffolds upon the sintering step, whichcan cause distortions and fracture of the models (Shanjaniet al., 2010). In addition, similarly to what was observedin scaffolds by SLS, residual non-bound material can befound in complex geometries and the surface can presenthigh roughness (Butscher et al., 2011).

5. Melt–extrusion-based techniques

Fused deposition modelling (FDM) is a commerciallyavailable AM technique, originally developed in 1992 forthe rapid manufacturing of prototypes for industry(Crump, 1992), which is based on the extrusion of a poly-meric filament through a heated nozzle (Figure 3d). Twoindependent extrusion nozzles can be used to deposit dif-ferent polymeric materials, one usually employed as asupport material and the other to produce the 3D object.

C. Mota et al.


The polymeric filament is supplied to the extrusion nozzleby means of drive wheels, and the polymer melt iscontinuously deposited on a construction platform. Thecomputer-controlled motion of the extrusion head andof the construction platform allows for the deposition ofthe molten filament in a predefined pattern (raster,contour and combination of both). After the fabricationof each layer, the platform moves downwards and thelayer-upon-layer process is performed until the 3D objectis obtained.

The possibility of producing TE scaffolds by FDM wasfirst explored by Hutmacher (2000), who reported theproduction of PCL or PCL–HA composite scaffolds withdifferent pore architectures, pore sizes and porosities. PCLscaffolds were shown to support the in vitro proliferation,differentiation and ECM production of primary humanfibroblasts and periosteal cells (Hutmacher et al., 2001).In vivo trials on a 2� 2 cm orbit defect created surgicallyon a Yorkshire pig model using PCL scaffolds prepared byFDM showed that the formation of new bone reached14.1% after 3months (Rohner et al., 2003). A clinical pilotstudy for cranioplasty on five patients employing PCL plugscaffolds fabricated by FDM showed that after 12monthsthe implants were well integrated in the surroundingcalvarial bone, with new bone filling the porous space(Schantz et al., 2006). After that, PCL scaffolds producedby FDM were approved by the FDA for craniofacial appli-cations (Osteopore International: www.osteopore.com.sg)and are currently marketed by Osteopore International inthe form of thin interwoven meshes (Osteomesh™) or 3Dimplants (Osteoplug™). A study reported on the implantationof Osteoplug™ scaffolds into burr holes of 12 patients treatedfor a chronic subdural haematoma. The scaffolds showedgood osteointegration into the surrounding calvarial boneand there were no adverse events in all patients, with ameanfollow-up of 16months (Low et al., 2009).

Biphasic constructs comprising a PCL cartilage scaffoldand a PCL–TCP osseous matrix produced by FDM wererecently investigated for the in vivo regeneration ofosteochondral defects (Ho et al., 2010). After implanta-tion into critical-sized osteochondral defects in pigs,mesenchymal stem cell (MSC)-seeded biphasic constructscoupled with an electrospun PCL–collagen membrane,acting as a barrier to prevent cell leakage, showed boneingrowth and remodelling as well as cartilaginous repair,with functional tissue restoration and low occurrence offibrocartilage. In addition, PCL/TCP scaffolds fabricatedby FDM were shown to be able to support in vitro humanMSC proliferation and osteogenic differentiation, andhuman MSC–scaffold constructs were implanted in nuderat critical-sized femoral defects, leading to cell survivalin the defect site for up to 3weeks post-transplantation,even though only 50% of the femoral defects respondedfavourably, as determined by new bone volume (Raiet al., 2010).

PCL scaffolds by FDM were also investigated incomparison with polyurethane sponges for the engi-neering of adipose tissue, showing in vivo angiogenesis,fibrous tissue formation and adipogenesis after 2 and

4weeks of implantation in nude mice (Wiggenhauseret al., 2012).

The FDM technique has given very promising preclinicaland clinical results, although its employment is so farlimited to a few thermoplastic polymers. Other melt–extrusion-based techniques have been developed in orderto process a wider range of polymeric materials in diverseforms (e.g. pellet, powder), reducing the quantities of rawmaterial used.

The 3D fibre deposition (3DF) technique, also called 3Dplotting, was originally developed by Landers andMülhaupt (2000) for the production of hydrogel scaffolds,but was subsequently employed for producing scaffoldsfrom a polymer melt, using a deposition system composedof a heating jacket wrapping the cartridge containing thepolymer. By applying a gas pressure, the molten polymeris forced through a nozzle and deposited selectively bymeans of an X–Y–Z moving arm. EnvisionTEC GmbH iscurrently marketing a 3DF system with the trademark3D-BioplotterW (Envisiontec: www.envisiontec.com/index.php?page=news&id=17).

Gloria et al. (2009) recently reviewed the potential andchallenges of 3DF to create multifunctional and tailor-madeTE scaffolds. Poly(ethylene oxide-terephthalate) and poly(butylene terephthalate) (PEOT/PBT) block co-polymersscaffolds produced by 3DF were investigated for cartilageregeneration purposes (Moroni et al., 2006). By changingscaffold porosity, fibre deposition and fibre orientation,the viscoelastic properties of this kind of scaffold could bemodulated to accomplish mechanical requirements fortailored tissue-engineered applications. Moreover, combin-ing an integrated non-woven electrospun microstructurewith a periodical 3DF macrostructure allowed a multi-scalePEOT/PBT scaffold with enhanced biological performancein terms of chondrocytes adhesion, proliferation, morphol-ogy and ECM production to be obtained (Moroni et al.,2008). Recently, PEOT/PBT scaffolds by 3DF were testedin vivo in comparison with analogous scaffolds fabricatedby compression moulding/salt leaching (Emans et al.,2012). Following 3weeks of in vitro culture in combinationwith allogenic chondrocytes, they were implanted intoosteochondral defects of skeletally mature rabbits showingimproved cartilage repair after 3months compared to com-pression-moulded scaffolds. 3D hybrid structures composedof a starch–PCL blend tubular scaffold, produced by 3DFand injected with gellan gum in the central hollow area,were investigated for spinal cord injury repair, showinggood in vitro compatibility with oligodendrocyte-like cells,olfactory ensheathing cells and Schwann cells (Silva et al.,2010, 2011). Moreover, preliminary in vivo studies,conducted in a hemisection rat spinal cord injury model,revealed that the hybrid scaffolds were well integratedwithin the injury and did not trigger any chronic inflamma-tory processes (Silva et al., 2011). A further study showedthat by changing the structure of this kind of scaffold interms of pore size and pore size gradients, it was possibleto vary scaffold mechanical properties as well as cellseeding efficiency and cell distribution within the scaffold(Sobral et al., 2011). Four types of PCL scaffold



http://www.osteopore.com.sg

http://www.envisiontec.com/index.php?page=news&id=17


architectures were recently investigated in vivo using a ratiliac crest defect model (Yilgor et al., 2012). New boneformation was higher on the scaffolds with higher porevolume for tissue ingrowth, and the healing of the bonedefects was further enhanced by combining the largestpore volume scaffolds with bone morphogenetic proteins.

Precision extrusion deposition (PED) is a melt–extrusionAM technique based on a screw extruder, enabling theprocessing of polymer pellets for the controlled depositionof a melt filament. PCL scaffolds with high precision at themicro-scale, as well as repeatability of the architecture,were developed by employing this technique (Wang et al.,2004; Shor et al., 2009). Osteoblast-seeded PCL scaffoldsshowed in vivo osseous ingrowth after subcutaneousimplantation in nude mice (Shor et al., 2009). Morevoer,the introduction of an assisting cooling device, whichincreases the working extrusion temperature up to 250�C,allowed the PED library of polymers to be expanded, asdemonstrated by producing 3D scaffolds made of PGA(melting point of 200�C) (Hamid et al., 2011).

Bioextruder is an innovative AM system comprising twodifferent deposition systems for simultaneously processingpolymer melts and solutions/suspensions to producemulti-material scaffolds incorporating cells and bioactiveagents (Mota et al., 2009). The melt–extrusion head iscomposed of a gas-pressurized container connected to ascrew extruder, allowing accurate control of polymer meltdeposition. This system was successfully employed toproduce PCL scaffolds with different pore sizes andarchitectures, showing that it is possible to control scaffoldporosity by varying different processing parameters,such as screw rotation velocity and deposition velocity(Figure 4g, h) (Domingos et al., 2009, 2010, 2011). TheBioextruder was recently combinedwith an electrospinningtechnique in order to develop a dual-scale structure,coupling the mechanical strength and structural reproduc-ibility of 3D microfilament constructs, fabricated by AM tothe advantage of electrospun ultrafine fibres in enhancingcell interaction with polymeric materials (Mota et al.,2011). PLGA electrospun fibres collected between adjacentPCL microfilaments offered a structural bridging for cellmobility and interaction, leading to cell colonization ofthe inter-filament gap (Figure 5).

Bioscaffolder™ (SysEng) is a commercially available AMmachine that comprises two different dispense headtypes: a syringe pump for gels, pastes or slurries; and amelt–extrusion head, composed of a gas-pressurizedreservoir coupled to an auger screw system for thermo-plastics (Ragaert et al., 2010). This machine was recentlyemployed to illustrate the lack of scaffold reproducibility inmelt–extrusion-based AM techniques when working withdegradation-sensitive polymers, such as PLA. PCL scaffoldsproduced by means of Bioscaffolder™ and embedded witha polyelectrolyte complex matrix of hyaluronic acid,methylated collagen and terpolymer showed in vitro higherseeding efficiency, more homogeneous cell distribution andhigher differentiation towards the osteoblastic phenotypeof human MSCs when compared to naked PCL scaffolds(Chen et al., 2011).

6. Solution/slurry extrusion-basedtechniques

Different AM techniques involving the processing of apolymeric solution or ceramic slurry have been developedin recent years (Figure 3e). They do not require the use ofhigh temperatures, thus avoiding concerns associatedwith temperature degradation of polymers and bioactiveagents, even though they often involve the use of organicsolvents that, if not completely removed, may compromisethe biocompatibility of the scaffolds and alter theincorporated bioactive factors. In addition to the previouslymentioned AM systems (BioplotterW, Bioscaffolder andBioextruder) that comprise several extrusion–depositionsystems, other approaches were developed for theproduction of scaffolds by processing solutions orslurries.

Pressure-assisted microsyringe (PAM) is an AM tech-nique employing a micropositioning system with apressure-activated microsyringe equipped with a fine-boreexit needle. It enables the controlled deposition of anextruded polymeric solution that, after solvent evaporation,leads to the formation of a predesigned microstructure(Vozzi et al., 2002, 2003). The possibility of producing 3Dstructures made of different biodegradable polymers, suchas PCL, PLLA or PLGA, by means of the PAM techniquehas been shown by different studies (Vozzi et al., 2002,2003). A piston-assisted microsyringe (PAM2) system wasrecently developed for low-shear stress extrusion of viscoushydrogel solutions containing cells, as demonstrated by astudy employing solutions of sodium alginate incorporatinghepatocytes (Tirella et al., 2011b).

Low-temperature deposition manufacturing (LDM) is atechnique exploiting a controlled cooling chamber thatallows a low temperature to be maintained throughoutthe deposition process of a polymeric slurry (Xiong et al.,2002). After the 3D structure is built up layer-by-layer, thesolvent is usually removed by freeze-drying. PLLA/TCPslurries were deposited using a computer-controlled LDMsystem to produce composite scaffolds for bone repair,achieving fairly good control over porosity and a micropo-rosity of 5mm in the structure strands, as consequence ofthe freeze-drying step. In vivo trials implanting the scaffoldson canine radius showed good biocompatibility and goodbone regeneration after 24weeks. A recent in vitro studyon composite scaffolds by LDM showed that PLGA–pearlscaffolds support the adhesion, proliferation and differenti-ation into osteoblasts of rabbit MSCs isolated from thefemoral crest (Xu et al., 2010). A similar AM systemequipped with a cryogenically cooled construction platformwas employed for the fabrication of natural polymerscaffolds (Kim et al., 2009, 2011). Using this technique,collagen scaffolds with controlled porous structure andlimited shrinkage were developed for skin TE (Kim et al.,2009). In order to increase structural stability, a hybridscaffold composed of an outer collagen shell and an inneralginate core was designed and successfully fabricated,showing enhanced mechanical properties and rapid

C. Mota et al.


vascularization when tested in vivo after implantation inmice as a dermal substitute (Kim et al., 2011).

Recent studies have investigated a new approach forthe fabrication of natural polymer scaffolds, involvingthe patterned extrusion of a photopolymerizable hydrogelink (composed of physically entangled polymer chainsdissolved in a monomer solution) under UV lightexposure. Using this technique, 3D hydrogel scaffolds weredeveloped by processing polymeric inks composed of poly(acrylamide) chains in a photopolymerizable acrylamidesolution (Barry et al., 2009), or poly(2-hydroxyethylmethacrylate) chains dissolved in a solution of its monomer(Hanson et al., 2011).

Robocasting, also referred to as direct-write assembly,is an AM technique allowing ceramic (e.g. b-TCP or HA)scaffolds to be fabricated with tailored geometry andporosity, using water-based, highly concentrated, colloidalsuspensions with minimal organic content that areextruded through a nozzle inside a non-wetting oil bath(Miranda et al., 2006, 2008). The compressive strength ofb-TCP scaffolds produced by robocasting was improved byinfiltration of melted PCL or PLA into the porous structure(Martinez-Vazquez et al., 2010). An alternative method ofproducing polymer/ceramic scaffolds with high inorganiccontent (70wt%) involved the processing of an HAsuspension in either PLA or PCL solution (Russias et al.,

2007). In this case, the high viscosity of the processed slurryallowed the utilization of the oil bath to be avoided and nosintering step was performed, due to the rapid evaporationof the solvent. Ceramic inks composed of HA, b-TCP orbiphasic HA/b-TCP suspension in PluronicW F-127 aqueoussolutions were used to produce scaffolds with differentmicroporosities and mechanical properties, depending onthe co-polymer content and the granulometry of theceramic powder (Franco et al., 2010). HA and b-TCPscaffolds produced by robocasting were also combined withBMP-2 growth factor to improve osteoinductivity, showingin vivo bone regeneration to be promoted after implantationin a pig model (Abarrategi et al., 2012). Suspensions of abioactive glass (6P53B) in PluronicW F-127 aqueoussolution was employed to fabricate scaffolds that, aftersintering, presented a compressive strength suitable for boneload-bearing applications (Figure 4i, j) (Fu et al., 2011).

7. Tissue and organ printing

Recent years have seen the rapid development of a newTE strategy, based on AM principles, commonly referredto as tissue or organ printing, which is aimed at the devel-opment of cell-laden constructs by employing automated

Figure 5. Development of dual-scale scaffolds by combining the Bioextruder system and the electrospinning technique (Mota et al.,2011): (a) bioextruder screw extrusion head for the fabrication of 3D PCL microfilament structures with a layer-by-layer approach;(b) electrospun PLGA ultrafine fibres collected on top of the PCL structures; (c) preosteoblast cells grown on the PCL structures;(d) preosteoblast cells grown on the dual-scale scaffolds



technologies and taking advantage of the self-organizingproperties of cells and tissues (Mironov et al., 2003).It involves the layer-by-layer deposition of cells andbioactive agents, usually in the form of self-assemblingtissue building blocks obtained by employing hydrogels toform biopaper or bioink (Figure 3f) (Mironov et al., 2007;Norotte et al., 2009). The aim is to deliver scaffoldingmaterials, living cells, nutrients, growth factors, drugs andother required chemicals with a precise control over thedeposition time, position and amount, in order to obtaincontrolled microenvironments for the development ofcustomized living tissue-engineered constructs (Bolandet al., 2006). An increasing number of research activitiesdedicated to this approach have led to the development oflaboratory-scale and commercially available manufacturingdevices based on different organ biofabrication technolo-gies, such as thermal or piezoelectric inkjet printing,3D printing by mechanical dispensing of cells and laser-assisted bioprinting (Guillemot et al., 2011; Mironovet al., 2011).

Thermal inkjet technologies have been explored for thedelivery of controlled volumes of delicate biologicalsystems to defined 3D locations (Boland et al., 2006,2007). As an example, aortic endothelial cells in combina-tion with a crosslinking calcium chloride solution weredeposited layer-by-layer, using a modified HP Deskjetprinter on top of a platform containing an alginate solution,proving that cells are able to attach and migrate in theresulting microchannel architecture (Boland et al., 2006).Using the same drop-on-demand process, micron-sizedfibrin channels were developed by employing humanmicrovascular endothelial cell suspensions mixed withthrombin as bioink and a fibrinogen solution as biopapersubstrate (Figure 6a, b) (Cui and Boland, 2009). A recentstudy involving the Olivetti BioJet system, which employsa thermal inkjet cartridge, investigated the velocity profileand mechanical load acting on a droplet during the cellprinting process (Tirella et al., 2011a). Moreover, by testingdifferent collection substrates, it was demonstrated that theimpact forces acting on the droplet during the depositionprocess, conditioning cell viability, are highly dependenton substrate stiffness.

A drop-on-demand inkjet printing system based on apiezoelectric membrane was investigated to deliversuspensions of human fibroblast cells, showing that theamplitude of the pulse used to excite the piezoelectricactuator has a small influence on cell survival rates (Saunderset al., 2008). However, cell agglomeration or sedimenta-tion inside the deposition cartridge affected the printingperformance.

Laser-assisted BioPrinting (LaBP) techniques haveemerged as an alternative approach to a cartridge-basedsystem for the production of tissue-engineered substitutes(Barron et al., 2004; Guillemot et al., 2010; Guillotin et al.,2010; Catros et al., 2011; Gruene et al., 2011; Koch et al.,2012). These techniques are based on laser pulses focusedonto a metallic absorption layer that evaporates at thefocal points, leading to the generation, on the subjacentlayer of cell-containing hydrogel precursor, of a jet

propelling towards a lower collector hydrogel slide.This technology is capable of positioning high celldensities (up to 108 cells/ml) and small volumes (downto a few hundred femtolitres) of cell suspensions withhigh resolution for the manufacture of 3D living con-structs with complex architecture (Ringeisen et al.,2006; Gruene et al., 2011; Koch et al., 2012). Forinstance, a recent study employed a LaBP system toreproduce the layered configuration of skin by printing,layer-by-layer, murine fibroblasts and human keratino-cytes embedded in collagen (Figure 6c, d) (Koch et al.,2012). Biological laser printing (BioLP™) was used toprint a variety of cell types, including osteosarcoma(Barron et al., 2004), olfactory ensheathing cells (Othonet al., 2008), carcinoma cells (Ringeisen et al., 2004),bovine aortic endothelial cells (BAECs) (Chen et al.,2006) and human umbilical vein smooth muscle celland human umbilical vein endothelial cells (HUVECs)(Wu and Ringeisen, 2010). In addition, HUVECs wererecently printed by BioLP™ onto hybrid biopapers thatcontained bilayers of PLGA and collagen type I to build3D vascularized tissues, resulting in good cell infiltrationand survival in the compound multilayer constructs(Pirlo et al., 2012).

As previously described, Bioscaffolder™ comprises,besides a melt–extrusion system, a pressure- or volume-controlled dispenser, allowing the accurate deposition ofpolymeric solutions and gels. Indeed, it was recentlyemployed for the processing of multipotent stromal cellsmixed with alginate to produce porous and solid hydrogelscaffolds, with cells homogeneously dispersed throughoutthe construct that, once implanted subcutaneously inimmunodeficient mice, resulted in ingrowth of vascular-ized tissue (Fedorovich et al., 2011a). A further studyshowed the suitability of Bioscaffolder™ to print intricateporous constructs containing two different cell types(endothelial progenitors and multipotent stromal cells)that retained heterogeneous cell organization afterin vivo implantation, leading to blood vessel formationin the endothelial progenitor cell-laden part and to boneformation in the multipotent stromal cell-laden part(Fedorovich et al., 2011b).

The recent explosion of interest in organ printing hasled to the development of other types of laboratory-available robotic systems for the production of cell-ladenliving constructs mimicking the structure of differenttissues, such as blood vessels (Norotte et al., 2009;Skardal et al., 2010), nerves (Francoise et al., 2012), skin(Lee et al., 2009; Koch et al., 2012) and bone (Fedorovichet al., 2011a, 2011b). However, as pointed out by Mironovet al. (2011), the use of robotic bioprinters alone is notsufficient for the development of large-scale industrialorgan biofabrication. In any case, progress in tissuespheroid biofabrication, the emergence of commercialbioprinters and the development of perfusion bioreactorssuitable for organ printing open new perspectives for thedesign of fully integrated organ biofabrication lines,necessary for the commercial translation of organprinting technology.

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8. Innovative AM equipmentTo overcome the limitations of the commercially availablesystems for TE scaffold production, several research groupsand specialized companies have developed customizedequipment that can meet different requirements in terms ofscaffold material, internal architecture and external shape.Some of the newly developed techniques are constantly beingadapted to offer the possibility of processing a wider range ofmaterials, improving the performance of the manufacturedstructures. One example of commercially available equipmentthat is continuously being improved is the 3D-BioplotterW sys-tem, which is nowadays in its fourth generation (Envisiontec:www.envisiontec.com/index.php?page=news&id=17).

Open-source AM systems have generated growing interestwithin the scientific community, offering adaptable tools fornew solutions in the manufacture of customized TE scaffolds.As an example, Fab@home is an open-source, open-architecture and low-cost rapid prototyping system, devel-oped originally for the production of end-user prototypes(Malone and Lipson, 2007). It is based onworldwide collabo-ration of users to develop customized AM systems withinnovative solutions, comprising new extrusion and/ordeposition heads that allow for the processing of noveland multiple materials in the course of scaffold fabrication.Recently, Cohen et al. (2011) used a Fab@home systemfor the development of alginate hydrogel scaffolds forcartilage replacements.

A fully automated bench-top manufacturing system,called BioCell Printing, for the integrated, continuous

and fully automated production and in vitro dynamicculture of tissue-engineering constructs was recentlydesigned (Bártolo et al., 2011). This system integrates fourphases (scaffold fabrication, sterilization, cell seeding anddynamic cell culture on the bioreactor) of the manufactureof tissue-engineered constructs in order to reduce risks ofcontamination that might occur in separated phases ofthe process and to reduce the manufacturing time.

A new research trend aims at direct printing in vivo andin situ of cells, biomaterials or their combinations. The firststudy using this approach involved laser printing of HAnanoparticles into mouse calvaria defects of critical size,resulting in an heterogeneous effect on bone formationfrom one sample to another (Keriquel et al., 2010). Anotherapproach recently developed comprises the printing in situof skin directly onto a wound by means of a cartridge-basedsystem (Binder et al., 2011). Robotic delivery systems,comprising an optical detector to collect data from thewounded region and computer-aided data processingsoftware to calculate the deposition pattern, weredeveloped and patented for in situ inkjet printing of cellsand biomaterials (Yoo et al., 2011).

9. Concluding remarks and futureperspectives

A wide range of AM techniques, based on various fabrica-tion strategies, have been proposed in the last 10 years for

Figure 6. Tissue-engineered constructs produced by means of different organ printing techniques: (a) scheme of thermal inkjettechnology and (b) confocal image of human microvascular endothelial cells inside a fibrin channel (Cui and Boland, 2009);(c) scheme of laser-assisted bioprinting (LaBP) technique for skin tissue engineering; and (d) histological image of murine fibroblastsand human keratinocytes embedded in collagen (Koch et al., 2012)




the development of TE scaffolds with different composi-tions and internal architectures. The increasing interestin these techniques is due to the possibility of achievinggood reproducibility and control over scaffold microstruc-ture and shape, and thus of developing customized tissue-engineered constructs that can meet specific requirementsin terms of pore size, geometry, interconnectivity, anatom-ical shape and size. In addition, in comparison withconventional scaffold fabrication techniques, they allowfor high automation of the manufacturing process,with a subsequent reduction of human intervention andenhancement of production rate.

Progress in AM technology has led to the developmentof laboratory prototype equipment that in some caseshas been finding commercial translation into advancedand versatile systems for TE scaffold production. Thenew frontier is the design of fully integrated AM linesfor the automation of the whole process necessary tofabricate ready-to-use tissue-engineered constructs on an

industrial scale. The idea is to integrate automated roboticdevices carrying out different stages in the developmentof engineered tissues, from preprocessing steps (medicalimaging, cell harvesting from patient, etc.) to tissuematuration by dynamic cell culture in bioreactors.

The rapidmanufacturing of customized tissue-engineeredconstructs tailored on specific patient requirements, besidesallowing fast production of large quantities of samples forhigh-throughput TE studies, can enhance clinical routineprocedures in terms of readily available products bettermeeting routine surgical needs.

Acknowledgements

This review on additive manufacturing techniques is intended tobe a follow-up to the scientific and technical inputs stemmingfrom projects financed by NoE Expertissues (EU-funded project,Grant No. NMP3–CT–2004–500283) and Hyanji Scaffold(EU-funded project, Grant No. PIRSES-GA-2008-230791).

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