3d bioprinting for musculoskeletal applications...

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3DBioprintingforMusculoskeletalapplications(Invited)

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DOI:10.2217/3dp-2017-0004

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191J. 3D Print. Med. (2017) 1(3), 191–211 ISSN 2059-475510.2217/3dp-2017-0004 © 2017 Future Medicine Ltd

part of

Review

This review focuses on developments in the field of bioprinting for musculoskeletal tissue engineering, along with discussion on the various approaches for bone, cartilage and connective tissue fabrication. All approaches (cell-laden, cell-free and a combination of both) aim to obtain complex, living tissues able to develop and mature, using the same fundamental technology. To date, co-printing of cell-laden and cell-free materials has been revealed to be the most promising approach for musculoskeletal applications because materials with good bioactivity and good mechanical strength can be combined within the same constructs. Bioprinting for musculoskeletal applications is a developing field, and detailed discussion on the current challenges and future perspectives is also presented in this review.

First draft submitted: 8 March 2017; Accepted for publication: 5 June 2017; Published online: 24 July 2017

Keywords:

Every 30 s, a patient dies from a condition that could be treated with organ replace-ment [1]. Organ transplantation has potential to be an efficient solution but is restricted due limited donor availability. Furthermore, organ transplantation requires complex sur-gical interventions and can lead to complica-tions such as organ dysfunction or rejection. Successful translation of tissue engineering and regenerative medicine research is key to alleviating the challenges in organ trans-plantation, but can also be applied to dis-ease modeling and drug discovery (Figure 1). More specifically, improved understanding of the biological architecture and natural repair processes in adult tissues could aid the challenging fabrication of de novo organs, for these applications. For cells to self-assemble into tissues, they need an environment in which cells can remain viable and are able to adhere and migrate. The most important fac-tors to consider are growth factors, nutrients, adhesion molecules, cells, materials and the

technologies applied to enhance the fabrica-tion process [2]. This review focuses on devel-opments in bioprinting for musculoskeletal tissue engineering, and provides discussion on the various approaches for bone, cartilage and connective tissue fabrication, along with current challenges and future perspectives.

Bioprinting & its role in musculoskeletal tissue fabricationThe musculoskeletal system (MSK) provides structural support for the body and comprises of vital tissue components such as bone, carti-lage, muscles, tendons and ligaments. When these tissues are damaged through injury, their repair remains challenging due to their limited regenerative potential.

Every year, over two million bone grafts are performed worldwide, due to diseases, sarcomas or trauma injuries [3]. In the USA, musculoskeletal injuries reach 32 million per year, of which, 45% are represented by ten-don, ligament and joint capsular i njuries [4].

3D bioprinting for musculoskeletal applications

Alexander Popov1, Sara Malferrari1 & Deepak M Kalaskar1

1

For reprint orders, please contact: [email protected]

192 J. 3D Print. Med. (2017) 1(3)

Figure 1. Translating tissue engineering and regenerative medicine research into healthcare applications for the future, including organ/tissue transplantation, disease modelling and drug discovery.

Cells

3D printing

Organ/tissuetransplantation

Bioreactor maturation

Disease modelling Drug discovery

Quality control

Biomaterials

Figure 2. 3D printing approaches for musculoskeletal tissue fabrication. The advantages and disadvantages of cell-free or cell-laden 3D printing.

mey

smes

memes

Cell-free printing

Advantages Disadvantages

sy c

Cell-laden printing

Advantagesy

cryce

DisadvantagesWe me

ys

mes

stresses

future science group

Review Popov, Malferrari & Kalaskar

Autografts are considered as the gold standard pro-cedure for treating small MSK defects due to their compatibility with patients. However, obtaining these grafts causes defects in secondary sites and can delay patient recovery time. Furthermore, autografts con-tribute additional trauma injury and are limited by size.

For bigger defects, allogenic or biomaterial grafts are used. However, the use of allogenic grafts has sev-eral disadvantages including: immune rejection, the necessity for multiple surgical procedures to remove donor material, cost and limited tissue regeneration

and revascularization potential [5]. Synthetic grafts can also inhibit vascularization and de novo tissue forma-tion [6]. Tissue engineering advances can provide solu-tions to these problems. Nevertheless, traditional tissue engineering approaches are slow and cannot be used for large scale production of biological matter with the required complexity.

Bioprinting promises fast, on demand, and auto-mated manufacturing of high resolution constructs. The process involves the use of 3D printing technolo-gies to deposit cells or biological factors into predefined shapes and sizes [7]. Bioprinting permits s tringent

193www.futuremedicine.comfuture science group

3D bioprinting for musculoskeletal applications ReviewTa

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194 J. 3D Print. Med. (2017) 1(3) future science group

Review Popov, Malferrari & KalaskarB

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/d

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ibu

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[17]

Tab

le 1

. Cel

l-fr

ee a

pp

roac

h f

or

bio

pri

nti

ng

of

mu

scu

losk

elet

al t

issu

es (

con

t.).


3D bioprinting for musculoskeletal applications Review

control on placement of cells within matrices and enables the arrangement of biological materials within composite, hierarchical structures and patterns. This promises new opportunities to fabricate reproduc-ible, patient-specific grafts with low risk of immune rejection. The most popular and promising bioprint-ing techniques include inkjet and extrusion printing. However, laser-assisted technologies are also in devel-opment [7]. Details on the various 3D bioprinting tech-niques are reported elsewhere [8].

To date, several research groups have bioprinted materials and cells for musculoskeletal applica-tions. The two most common approaches used for 3D bioprinting include the printing of cell-free and cell-laden materials. Tables 1 & 2 outline how bone, cartilage, muscle, tendon and ligament tissues have been fabricated using these approaches. Figure 2 s ummarizes the benefits of each approach.

The current literature suggests that synthetic mate-rials are more widespread for cell-free printing, while natural polymers have been commonly combined with cells, prior to extrusion. To evaluate the most popular approaches and materials used for bioprinting of musculoskeletal tissues, we reviewed the l iterature p ublished in this area during the last 15 years. Figure 3 summarizes the various materials used for 3D print-ing using cell-free and cell-laden approaches in MSK tissue engineering. The data suggest that around 84% of the materials used for cell-free printing are synthetic. This is primarily because these materials provide the strong mechanical properties required for musculoskeletal applications. The remaining 16% of articles show feasibility of this cell-free approach using natural materials, such as collagen and alginate, and this can be due to their higher biocompatibility compared with synthetic materials.

One good example of cell-free printing for bone repair focused on composite scaffold printing, where polycaprolactone-hydroxyapatite-carbon nanotubes were printed with pores in the range of 450–700 μm (Figure 4) [15]. This work highlights that cell-free bio-logical materials can be classed as bioinks. Results show that a 4 MPa compressive strength was obtained and this is analogous to trabecular bone. The com-posite scaffolds also exhibited enhanced cell adhe-sion and improved hydroxyapatite bioactivity, when seeded with MG63 osteoblast-like cells. Nevertheless, the materials were distinct from native bone due to the presence of PCL and carbon nanotubes, and the cells were seeded in a traditional engineering approach. This led to uneven cell seeding/distribution.

In the literature, naturally derived biomateri-als, such as alginate [14,20], collagen [14], gelatin and fibrin [25], have successfully promoted cell adhesion, B

iom

ater

ials

Cel

lsPr

inti

ng

te

chn

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licat

ion

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ruct

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pro

per

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van

tag

es (

A)/

dis

adva

nta

ges

(D

)Re

f.

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llag

en +

al

gin

ate

+ si

lica

Mo

use

pre

-o

steo

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sts

(MC

3T3-

E1)

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te

mp

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ure

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iop

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er

Bo

ne

Mu

lti-

laye

red

cyl

ind

rica

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(32

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h

mes

h-l

ike

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stru

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re. H

igh

ly p

oro

us

(>78

%)

wit

h 4

68–4

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m

aver

age

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re s

ize

Ten

sile

Yo

un

g’s

mo

du

lus:

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6 ±

0.19

MPa

M

ax. t

ensi

le s

tren

gth

: 0.1

2 ±

0.03

M

Pa

Co

mp

ress

ive

You

ng

’s

mo

du

lus:

∼0.

2–0.

3 M

Pa

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ioco

mp

atib

ility

, ost

eo-

ind

uct

ion

an

d p

rod

uct

ion

o

f b

on

e-lik

e H

A. S

ilica

im

pro

ved

mec

han

ical

p

rop

erti

es c

om

par

ed

wit

h c

olla

gen

+ a

lgin

ate

hyd

rog

els

on

ly

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-ste

p s

caff

old

fa

bri

cati

on

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d c

ell

seed

ing

, wit

h >

7-d

ay

coat

ing

pro

cess

th

at c

an

cau

se b

lock

ed p

ore

s

[14]

Silic

on

-do

ped

n

ano

crys

talli

ne

HA

+ P

CL

+ C

NT

Hu

man

o

steo

sarc

om

a (M

G 6

3)

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mat

ic

Envi

sio

nTEC

3D

B

iop

lott

er®

Bo

ne

Mu

lti-

laye

red

latt

ice.

7

laye

rs w

ith

6 m

m

dia

met

er a

nd

3 m

m

hei

gh

t. In

terc

on

nec

ted

sq

uar

e 45

0–7

00 μ

m p

ore

s

Co

mp

ress

ive

stre

ng

th: ∼

4 M

Pa

Co

mp

ress

ive

elas

tic

mo

du

lus:

50

MPa

A: C

NTs

imp

rove

cel

l at

tach

men

t. 2

% C

NT

scaf

fold

s im

pro

ve

mec

han

ical

pro

per

ties

an

d

elec

tric

al c

on

du

ctiv

ity

D: s

caff

old

s lo

aded

wit

h

mo

re t

han

2%

CN

Ts

dec

reas

e co

mp

ress

ive

resi

stan

ce a

nd

po

rosi

ty

(40%

)

[15]

Gel

MA

MG

63

ost

eob

last

-lik

e ce

lls

Prim

ary

NH

Ost

Cu

sto

miz

ed

bio

pri

nte

rB

on

ePo

res

size

400

μm

, th

ickn

ess

750 μm

H

ydro

gel

wit

h 8

% G

elM

A

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ore

cro

sslin

kin

g: s

tora

ge

mo

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100

Pa

Aft

er U

V c

ross

linki

ng

: sto

rag

e m

od

ulu

s 10

00 P

a

A: s

tora

ge

mo

du

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per

mit

s p

rin

tin

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f th

e h

ydro

gel

bef

ore

cr

oss

linki

ng

, an

d U

V-

cro

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re

suit

able

mec

han

ical

p

rop

erti

es t

o s

tim

ula

te

ost

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last

s p

rolif

erat

ion

. G

elM

A h

ydro

gel

has

su

cces

sfu

lly b

een

use

d t

o

coat

tit

aniu

m

D: l

ow

cel

l via

bili

ty

[16]

PCL

+ PL

GA

+

du

ck b

eak

New

Zea

lan

d

Wh

ite

rab

bit

in

viv

o

stu

dy

wit

h 5

m

m c

riti

cal

def

ects

Bio

pri

nti

ng

: m

ult

i hea

d

pn

eum

atic

sy

rin

ge

dis

pen

ser

Bo

ne

3 ×

3 ×

20 m

m o

blo

ng

sc

affo

lds

wit

h 7

7.3%

p

oro

sity

an

d 2

.787

μm

p

ore

s

Co

mp

ress

ive

stre

ng

th: 1

7 M

PaA

: pro

mo

tes

rep

air

and

d

e n

ovo

bo

ne

form

atio

n.

Hig

h c

om

pre

ssiv

e st

ren

gth

co

mp

ared

wit

h P

CL/

PLG

A

imp

lan

ted

sca

ffo

lds

D: i

rreg

ula

r sc

affo

ld

shap

e an

d p

ore

str

uct

ure

/d

istr

ibu

tio

n

[17]

Tab

le 1

. Cel

l-fr

ee a

pp

roac

h f

or

bio

pri

nti

ng

of

mu

scu

losk

elet

al t

issu

es (

con

t.).

Bio

mat

eria

lsC

ells

Prin

tin

g

tech

niq

ue

Ap

plic

atio

nC

on

stru

ct m

orp

ho

log

yM

ech

anic

al p

rop

erti

esA

dva

nta

ges

(A

)/d

isad

van

tag

es (

D)

Ref.

PLA

+H

AIn

viv

o

imp

lan

tati

on

o

f M

SCs

fro

m

New

Zea

lan

d

rab

bit

s

Des

kto

p

3D p

rin

ter

(Do

t G

o 3

D

Tech

no

log

y C

orp

ora

tio

n,

Xia

ng

tan

, C

hin

a)

An

teri

or

cru

ciat

e lig

amen

t

Scre

w-l

ike

scaf

fold

, 10

× 2.

1 ×

2.1

mm

Po

res ∼2

90 μ

m

N/A

A:in

viv

o w

ork

sh

ow

ed

go

od

bo

ne

/gra

ft in

terf

ace

and

su

cces

sfu

l ten

do

n

hea

ling

wit

hin

bo

ne

tun

nel

D

: no

mec

han

ical

tes

t an

d

scaf

fold

no

t re

pre

sen

tati

ve

of

ph

ysio

log

ical

en

viro

nm

ent

[18]

196 J. 3D Print. Med. (2017) 1(3)

Figure 3. 3D bioprinting for musculoskeletal applications. Percentage of various materials used in (A) cell-free and (B) cell-laden printing for musculoskeletal applications. The data are based on articles published in the last 15 years using the search terms ‘3D printing’, ‘bioprinting’ and ‘bioink’ associated with ‘bone’, ‘cartilage’, ‘osteochondral’, ‘muscle’, ‘tendon’ and ‘ligament’.

Silica6%

Collagen5%

GelMA6% PLGA

17%

PLA11%

HA22%

PCL17%

β-TCP5%

Alginate11%

Cell-free printing

Cell-laden printing

Cellulose9%

Polyurethane3%

PLGA3% HA

3%

PCL21%

Alginate25%

Collagen9%

Agarose3%

Hyaluronic acid9%

Gelatine6%

PEGDMA3%

Fibrin3%

dECM3%



proliferation and osteochondral differentiation with both cell-free and cell-laden printing approaches. However, natural materials were the most com-mon materials used for cell-laden printing (70% of total reported in literature, Figure 3B) for musculo-skeletal applications. This is predominantly because of their capacity to form gels that can support cell e ncapsulation and survival during the 3D printing process.

Due to their poor mechanical properties, some of these natural biomaterials have been co-printed with other synthetic polymers for musculoskeletal applica-tions. This can be realized by using multi-tool printing, which requires special modifications to printers such as incorporation of additional print heads or extruders. Daly et al. used multi-tool printing to produce a mechan-ically reinforced cartilaginous template mimicking the geometry of a vertebral disk [26]. This is an exciting study



ble

2. C

ell-

lad

en a

pp

roac

h f

or

bio

pri

nti

ng

of

mu

scu

losk

elet

al t

issu

es.

Bio

mat

eria

lsC

ells

Prin

tin

g t

ech

niq

ue

Ap

plic

atio

nC

on

stru

ct

mo

rph

olo

gy

Mat

eria

l p

rop

erti

esA

dva

nta

ges

/dis

adva

nta

ges

Ref

.

Lutr

ol F

127

+ M

atri

gel

®

+ al

gin

ate

+ m

eth

ylce

llulo

se

+ ag

aro

se

Go

at b

on

e m

arro

w

stro

mal

cel

lsPn

eum

atic

En

visi

onT

EC 3

D

Bio

plo

tter

®

Bo

ne

20 ×

20

mm

re

ctan

gu

lar

3D

scaf

fold

s w

ith

300

μm

sp

acin

g b

etw

een

fi

ber

s an

d 1

50 μ

m

laye

r th

ickn

ess

N/A

A: c

ell v

iab

ility

no

t af

fect

ed

by

pri

nte

r n

ozzl

e. M

ater

ials

p

rom

ote

d c

ell p

rolif

erat

ion

an

d

dif

fere

nti

atio

n in

to o

steo

gen

ic

linea

ge

D: h

ydro

gel

s st

iffn

ess

lead

ing

to

fu

sio

n o

f ad

jace

nt

laye

rs, a

nd

no

tr

ansv

erse

po

res

[19]

Alg

inat

eH

um

an a

rtic

ula

r ch

on

dro

cyte

s +

hM

SCs

Bio

Scaf

fold

er™

p

neu

mat

ic s

yste

mO

steo

cho

nd

ral

Ten

-lay

er

rect

ang

ula

r 3D

sc

affo

lds

(10

× 10

m

m)

wit

h s

pac

ing

b

etw

een

fib

ers

of

0.8

–2.5

mm

. 100

μm

la

yer

thic

knes

s

Vis

cosi

ty: 0

.5–

1000

Pa.

sA

: co

ntr

olle

d c

ell d

istr

ibu

tio

n/

enca

psu

lati

on

in h

ydro

gel

s.

Co

mb

ine

mu

ltip

le c

ell t

ypes

D

: po

or

mec

han

ical

str

eng

th o

f al

gin

ate

and

fu

sed

tra

nsv

ersa

l p

ore

s

[20]

Cel

l-la

den

al

gin

ate

surr

ou

nd

ing

ce

ll-fr

ee P

CL

(cel

l-fr

ee)

Hu

man

ch

on

dro

cyte

s (C

20A

4)B

ioSc

affo

lder

™

pn

eum

atic

sys

tem

Har

d t

issu

e6

× 60

× 2

mm

re

ctan

gu

lar

latt

ice

scaf

fold

s w

ith

2 μ

m

fib

er s

pac

ing

You

ng

’s m

od

ulu

s:

∼6.5

MPa

A: P

CL

imp

rove

s m

ech

anic

al

pro

per

ties

of

alg

inat

e D

: hig

h d

epo

siti

on

tem

per

atu

res

of

PCL

are

det

rim

enta

l to

cel

l vi

abili

ty in

alg

inat

e

[21]

PCL

+ PL

GA

+

atel

oco

llag

enM

ou

se p

re-

ost

eob

last

s (M

C3T

3-E1

)

Bio

pri

nti

ng

: m

ult

ihea

d p

neu

mat

ic

syri

ng

e d

isp

ense

r

Het

ero

gen

eou

s ti

ssu

eH

ybri

d s

caff

old

wit

h

alte

rnat

ed la

yers

of

syn

thet

ic a

nd

nat

ura

l m

ater

ials

, wit

h 4

00

μm fi

ber

s

N/A

A: s

caff

old

pro

mo

tes

cell

pro

lifer

atio

n a

nd

go

od

via

bili

ty

D: n

o c

on

sid

erat

ion

of

mec

han

ical

p

rop

erti

es

[22]

PCL

+ h

yalu

ron

ic a

cid

+

atel

oco

llag

en

Hu

man

mes

ench

ymal

st

rom

al in

rab

bit

kn

eeB

iop

rin

tin

g: m

ult

i h

ead

pn

eum

atic

sy

rin

ge

dis

pen

ser

Ost

eoch

on

dra

lH

ydro

gel

s w

ith

250

an

d 5

00 μ

m p

ore

s,

bet

wee

n 2

50 μ

m P

CL

fib

ers.

5 m

m r

abb

it

knee

def

ect

fille

d

N/A

A: m

ult

ilaye

red

co

nst

ruct

s w

ith

ou

t ch

emic

al o

r p

hys

ical

cr

oss

linki

ng

. PC

L al

low

ed c

ell

rich

hyd

rog

els

wit

h c

on

tro

lled

st

ruct

ure

D

: no

co

nsi

der

atio

n o

f m

ech

anic

al

pro

per

ties

[23]


Review Popov, Malferrari & KalaskarTa

ble

2. C

ell-

lad

en a

pp

roac

h f

or

bio

pri

nti

ng

of

mu

scu

losk

elet

al t

issu

es (

con

t.).

Bio

mat

eria

lsC

ells

Prin

tin

g t

ech

niq

ue

Ap

plic

atio

nC

on

stru

ct

mo

rph

olo

gy

Mat

eria

l p

rop

erti

esA

dva

nta

ges

/dis

adva

nta

ges

Ref

.

PEG

DM

AH

um

an a

rtic

ula

r ch

on

dro

cyte

sTh

erm

al in

kjet

p

rin

ter:

Hew

lett

-Pa

ckar

d D

eskj

et 5

00

Car

tila

ge

Cylin

dri

cal

ost

eoch

on

dra

l plu

gs,

4

mm

in d

iam

eter

an

d 2

mm

in d

epth

Co

mp

ress

ive

mo

du

lus:

321

.06

± 43

.99

kPa

Swel

ling

rat

io:

6.10

–11.

80%

A: s

imu

ltan

eou

s p

ho

to-

po

lym

eriz

atio

n d

uri

ng

3D

pri

nti

ng

to

mai

nta

in p

reci

se c

ell p

osi

tio

n

du

rin

g la

yer-

by-

laye

r as

sem

bly

. In

teg

rate

d la

yers

dec

reas

e d

elam

inat

ion

ris

k. C

om

pre

ssiv

e m

od

ulu

s co

mp

arab

le t

o n

ativ

e h

um

an a

rtic

ula

r ca

rtila

ge.

B

ioco

mp

atib

ility

an

d p

rom

oti

on

o

f ch

on

dro

cyte

gro

wth

D

: co

mp

ress

ive

mo

du

le lo

wer

th

an n

on

pri

nte

d P

EGD

MA

du

e to

th

erm

al d

egra

dat

ion

[24]

Gel

atin

-fib

rin

m

atri

xh

MSC

s +

hU

VEC

s +

hu

man

neo

nat

al

der

mal

fib

rob

last

s

Aer

ote

ch A

GB

10,

000

pn

eum

atic

syr

ing

e d

isp

ense

r

Bo

ne

Hyd

rog

el in

a 3

D

per

fusi

on

ch

ip (

725

× 65

0 ×

125

mm

)

Vis

cosi

ty: ≤

1000

Pa

.s

Shea

r el

asti

c m

od

ulu

s:

1–10

,000

Pa

Plat

eau

mo

du

lus:

30

0–5

000

Pa

Yie

ld s

tres

s:

700

–900

0 Pa

A: p

rolif

erat

ion

an

d

dif

fere

nti

atio

n in

to o

steo

gen

ic

linea

ge

aro

un

d v

esse

ls p

erfu

sed

w

ith

ost

eog

enic

med

ium

. Cel

ls

surv

ive

for

mo

re t

han

6 w

eeks

. Th

ick

(>1

cm)

vasc

ula

rize

d

con

stru

ct

D: w

eak

mec

han

ical

pro

per

ties

of

the

con

stru

ct d

ue

to g

elat

in a

nd

fi

bri

n p

rop

erti

es (

no

t as

sess

ed)

[25]

Alg

inat

e +

PCL

Porc

ine

MSC

s3D

Dis

cove

ry m

ult

i-h

ead

pn

eum

atic

b

iop

rin

tin

g s

yste

m

Bo

ne

Hu

man

ver

teb

rae-

like

stru

ctu

res

wit

h o

rth

og

on

al,

1 m

m P

CL

fib

ers

in

cell-

lad

en a

lgin

ate

hyd

rog

el

Co

mp

ress

ive

mo

du

lus:

160

0 ±

100

Pa

A: P

CL

rein

forc

ed m

ech

anic

al

pro

per

ties

of

alg

inat

e. C

on

stru

ct

pro

mo

ted

en

do

cho

nd

ral

oss

ifica

tio

n a

nd

vas

cula

riza

tio

n

po

stim

pla

nta

tio

n

D: c

o-p

rin

tin

g o

f M

SC-l

aden

al

gin

ate

wit

h P

CL

no

t p

oss

ible

in

sm

alle

r d

iam

eter

co

nst

ruct

s (<

6 m

m).

Alg

inat

e an

d P

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201www.futuremedicine.com

Figure 4. Structural properties of PCL-hydroxyapatite composites strengthened with carbon nanotubes for bone repair. (A) Technical drawing, (B) 3D simulation and (C) printed 3D scaffold with square pores. Reproduced with permission from [15].

Figure 5. Bioprinting of vertebrae-shaped bioinks with enhanced mechanical properties for bone regeneration. (A) Multi-tool 3D printing was applied to produce a composite vertebrae structure by depositing PCL filaments followed by an MSC-laden bioink (RGD-alginate). (B) MicroCT analysis to illustrate the distribution of the bioink and PCL. (C) Live/dead images of MSCs within the bioprinted vertebrae. MSC: Mesenchymal stem cell; PCL: Polycaprolactone; RGD: Arginylglycylaspartic acid. Reproduced with permission from [26].

2. PCL deposition 3. Bioink/MSC deposition1. Vertebral bodySTL

4. Composite vertebrae



showing the potential to mature bioprinted constructs in vivo. In this approach developmental precursors to an adult organ were bioprinted and the engineered con-struct functioned as a template for subsequent organo-genesis in vivo. This was achieved by printing a polycap-rolactone (PCL) template, f ollowed by the deposition of

a arginylglycylaspartic acid (RGD)-alginate hydrogel laden with adult stem cells, as shown in Figure 5. When implanted into a mouse model, the resultant bioprinted construct supported the development of vascularized bone containing trabecular-like endochondral bone with a s upporting marrow structure.

202 J. 3D Print. Med. (2017) 1(3)

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In another example using multi-tool printing, the vascularized bone was engineered when polylactic acid (PLA) was deposited with fused deposition modelling (FDM) printing technology, while gelatine methac-ryloyl (GelMA) containing BMP and VEGF was co-printed using stereolithography (SLA) [34]. This study demonstrated that the dual 3D printed constructs pro-vided a hierarchically biomimetic bone-like structure, with multiphasic characteristics and potential for vas-cularized bone regeneration, as shown in Figure 6. This is a noteworthy approach to produce complex tissue structures in the lab. However, it is important to strin-gently assess the functionality of this vascularized bone and make quantitative comparisons with native tissues.

In addition to this work, interesting results have been obtained using synthetic materials. Polyethylene glycol dimethyl acrylate was successfully inkjet printed with human chondrocytes for the repair of small osteo-chondral defects [24]. In another study, bone and car-tilage tissues were fabricated through inkjet printing of human mesenchymal stem cells and simultaneous deposition and photo-crosslinking of polyethylene gly-col dimethyl acrylate [35]. Others have achieved mini-mal print-head clogging by printing acrylate peptides and polyethylene glycol hydrogels with human mesen-chymal stem cells to promote robust bone and cartilage formation [36]. Porous bioactive glass/alginate compos-ite scaffolds have also been fabricated for bone tissue engineering using 3D printing [13].

As a prospective treatment for cartilage lesions, recent study reported use of 3D bioprinting approach to form cartilage mimetics using a nanofibrillated cel-lulose and alginate-based composite bioink seeded with human-derived induced pluripotent stem cells and human chondrocytes [30]. The bioprinted con-structs could maintain pluripotency initially, and after 5 weeks, hyaline-like cartilaginous tissue with collagen type II expression, lacking tumorigenic OCT4 expres-sion was observed. Furthermore, a significant increase in cell number within the cartilaginous tissue was detected. This study combines 3D printing and stem cell technology to generate viable tissues for clinical applications.

While the majority of studies focus on bone and car-tilage regeneration, recent work has shown progress in the bioprinting of other musculoskeletal tissues such as muscles and tendons. An integrated tissue-organ printer was used for the fabrication of skeletal muscle structures. The approach was based on the printing of well-defined PCL patterns for directional alignment of the muscle cells, as shown in Figure 7. At the same time, cells were deposited using a mixture of hydrogels (gela-tin, hyaluronic acid and fibrinogen), which were loaded with mouse myoblasts cells. Results showed good cell

203www.futuremedicine.com

Figure 7. Bioprinting of skeletal muscle and implantation in vivo. (A) and (B) scaffold design; (C) scaffold fabrication; cell alignment with PCL (D) and without PCL (E); (F) live/dead assay: green cells are alive and red cells are dead; (G) immunofluorescent staining for myosin heavy chain of the 3D printed muscle after 7D differentiation. The encapsulated myoblasts aligned along the longitudinal direction of the fiber structure; (H) schematic of the ectopic implanted scaffold in vivo; (I) implanted scaffold next to the CPN and (J) immunostaining for desmin. CPN: Common peroneal nerve; PCL: Polycaprolactone. Reproduced with permission from [31] © Nature Publishing Group.

Commonperonealnerve(CPN)

Bioprintedmuscle

Gluteusmuscle

I J

Cell-ladenhydrogel pattern

Cross-sectional view

PCL pillar

Cell-laden hydrogel

After removingPluronic F-127

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viability, and alignment along the PCL pillars/patterns and muscle-like structures were observed after 7 days. When the constructs were implanted in vivo, they inte-grated with the common peroneal nerve after 2 weeks and the muscle was seen to respond to electrical stim-uli [31]. Even though bioprinting examples in this area are limited, this study provides good evidence that 3D printing can be used for the development of various fibrous tissues (muscle, tendon and ligament) where cellular alignment is a key r equirement [37].

While various materials have been used as bioinks for printing cell-free and cell-laden constructs, cells alone in the form of tissue spheroids have also been

i nvestigated for 3D bioprinting. Printed cells have a fluid nature and over time, they fuse together to form more complex cell aggregates that can potentially lead to tissue formation [38]. In the literature, tissue spheroids have already been successfully used for cartilage tissue engineering [39–41]. However, successful production of constructs using tissue spheroids is still in its infancy and focus needs to be applied on utilizing 3D printing technologies to help with scale-up, r eproducibility and formation of more complex structures [42,43].

Breakdown of the materials used as bioinks for bone and cartilage bioprinting in the last 15 years show some interesting trends. The majority of all the cell-free

204 J. 3D Print. Med. (2017) 1(3)

Figure 8. Characteristic requirements for bioinks in musculoskeletal tissue fabrication.

Rheological properties

Printability Biocompatibility

Vascularisation capacity

Cell differentiation andtissue development

Biodegradation

Cell adhesion,proliferation andmigration

Tissue-specific ECMmimicry

ResolutionMalleability

ReproducibilityCost

Fabrication timeResistance to printingstresses (pressure,temperature and shearforces)

Physico-chemical

Fabrication Biological

Crosslinking/gelationtimeSwelling/contractilityPorosityDegradation ratepH/ionic sensitivity

Mechanical strength/stiffness

Bioinks



approaches used materials such as PCL, β-tricalcium phosphate or hydroxyapatite (Figure 3). In contrast, alginate was the most popular material for cell-laden bioinks (25%) and this was followed by PCL (21%) and collagen (9%). Alginate was applied due to its good printability, while PCL is a biocompatible mechanical strength enhancer of the cell-laden hydrogels. Notably, there is greater variety in the materials used for cell-laden printing than cell-free printing.

3D printing has been successfully applied in a variety of ways to address the growing demand for more robust musculoskeletal therapies. Nevertheless, the use of the technology for medical purposes is still in its infancy. There is need for further research and development in both 3D printing technology and bio-ink formulations for successful translation of this t echnology in future.

Musculoskeletal bioinks & their characteristicsBioinks are an integral part of the bioprinting process. Most frequently, they are defined as hydrogel materials used for the encapsulation of cells in 3D bioprinting [44]. However, this definition is very limited and there are a number of examples where b iological m aterials wi thout

cells are also termed as bioinks [45]. Opinion on the exact definition of bioink remains divided. From the literature, it is clear that cell-only [46], cells with sup-porting materials (both synthetic and natural hydro-gels) [26] and biomolecules without cells (BMP2) [47] are also referred to as bioinks.

In the musculoskeletal context, the scenario is even more complex. As seen from the literature review in the previous section, musculoskeletal tissues have been bioprinted using three approaches: cell-free, cell-laden and combination of both approaches (i.e., multimode printing of synthetic polymers, along with encapsu-lated cells). The definition of bioinks becomes even hazier as the commonly used definition of encapsu-lated cells within material becomes very limited in its scope and application. We anticipate that as bioprint-ing research advances through the development of both hardware (3D printers) and novel materials that support this process, the need for an accurate and more inclusive definition will become apparent. Here we will focus on the various bioink used for musculoskeletal applications as reported in literature and define the requirements for the fabrication of functional MSK tis-sues. Figure 8 shows physical–chemical, biological and



Table 3. Summary of the most popular hydrogels used for musculoskeletal bioprinting.

Material Description Tissue Ref.

Natural hydrogels

Alginate Polysaccharide derived from seaweed, which can be ionically crosslinked with CaCl2. A: fast gelation, general ease of use and low cost. D: low swelling properties can limit cell survival and growth in the long term, and weak mechanical properties (compressive modulus ∼10 kPa). CaCl2 crosslinker can be cytotoxic at high concentrations

Bone [26,62]

Chitosan Polysaccharide derived from chitin and most commonly obtained from the exoskeletons of crustaceans. Commonly crosslinked using genipin or glutaraldehyde. A: good biocompatibility, biodegradability, anti-inflammatory/antibacterial properties and good printability. Structure similar to GAGs in cartilage. D: slow gelation and weak mechanical properties (Young’s modulus in compressive mode for noncrosslinked and genipin crosslinked films is 38.7 and 87.3 kPa). Glutaraldehyde crosslinking is cytotoxic

Bone and cartilage

[63]

Collagen Protein composed of glycine, praline and arginine to form tropocollagen fibers of diameters ranging from 50 to 200 nm. A: good swelling properties and biocompatibility. D: weak mechanical strength (mean peak stress: 0.76 MPa), poor printability and expensive

Bone and cartilage

[64]

Fibrin Fibrin is a nonglobular protein present in the blood produced during blood clotting. Can be enzymatically crosslinked. A: good biocompatibility, swelling and gelation properties. Good printability. D: weak mechanical properties and expensive

Bone [65]

Gelatin Protein obtained from hydrolyzed collagen, which can be crosslinked using temperature and enzymes. A: low cost, biocompatible, biodegradable with high-cell adhesion. D: poor printability. Often found coupled to MA to form GelMA (crosslinked with UV light and harmful to cells), which has significantly increased mechanical properties and improved printability compared with gelatin

Bone and cartilage

[66,67]

HYA Polysaccharide and major component of ECM. (Photo) chemical crosslinking. A: good biocompatibility, good swelling ratio (0–45) fast degradation rates (100–0% residual mass in 8 days). D: weak mechanical properties (storage modulus 400–1000 Pa, loss modulus 3–30 Pa) and limited printability due to shear-thinning

Bone [68]

dECM Tissue decellularized ECM can be obtained using chemical, physical and biological treatments. A: representative of natural ECM environment, tissue-specific, guides for stem cell differentiation and good biocompatibility. D: nonhomogeneous cell seeding and immune reactions if cellular components remain. Decellularization treatments can damage natural ECM. Poor mechanical properties: max storage modulus of 300 and 0–20 Pa loss modulus

Bone Skeletal muscle

[69] [32]

Synthetic hydrogels

PEG Synthetic polymer. Commonly crosslinked via chain-growth and step-growth polymerization, but can also be crosslinked using radiation and other chemical/physical methods. A: US FDA approved and does not trigger immunological responses. The material is soluble in water and organic solvents and has low protein adhesion properties. Good diffusion of nutrients and oxygen, and cell migration. When coupled to MA, it shows high swelling properties (swelling ratio from 37.88 to 100.93%) and good biocompatibility. Compressive modulus of 30–65 kPa and 1.63–6.99 cP viscosity. Good printability. D: often lack bioactive molecules

Bone [36]

Self-assembling peptides

Self-assembling peptide-based gels. A: versatile properties that can be easily tailored by adjusting chemicals and physical parameters. Good extracellular matrix mimicry, while biocompatible and biodegradable. Bioactive molecules can be incorporated. D: peptide gels are degraded by cellular proteolytic enzymes to cause shrinkage (20% of hydrogel volume in 12 days). Consequently, the mechanical properties become weaker (decrease from 50 to 10 Pa in 12 days)

Bone [70]

A D



fabrication requirements for musculoskeletal bioinks. Most of these requirements are similar to soft tissue bioprinting; however, they become specific for mus-culoskeletal applications when additional mechanical stiffness and rigidity is required to fabricate structur-ally competent tissues.

In the body, cells are found in highly organized environments, which are rich in water, nutrients and growth factors [48]. The review paper by Murphy pro-vides an excellent overview Due to their significant water content, hydrogels have been identified as a primary material for bioprinting. Additionally, their hydrophilic nature allows hydrogels to retain large vol-umes of water without preventing a variety of cross-linking methods to be applied during fabrication of 3D networks [49]. In addition to this, hydrogels can be formulated to respond to various external stimuli such as temperature, electric or magnetic fields, light, pressure and sound vibrations before, during or after printing process [50]. Chemical factors including pH, solvent composition, ionic strength and molecular species also affect hydrogel properties. Therefore, a good understanding of these parameters on printabil-ity, stability in both in vitro and in vivo environments becomes essential.

Materials for bioprinting must be biocompatible and mimic natural cellular or tissue environment [51,52]. Specifically, materials used for cell encapsulation must mimic the natural environment of cells and it has been demonstrated that hydrogels based on extracellular matrix components permit this [53,54].

In terms of fabrication, printing materials should exhibit good printability and sufficient mechanical properties for cellular support and maintenance of the 3D structures [51]. For example, since highly viscous hydrogels are prone to clogging phenomena in the nozzles of extrusion-based printers, shear-thinning of some hydrogels, such as hyaluronic acid and peptide gels, can be advantageous [55,56]. However, it is impor-tant to adapt these hydrogels so they are able to ‘self-heal’ and maintain their printed structure once depos-ited [57]. Gelation time, along with the capacity to respond to physiological shear, tensile and compressive stresses, are other key parameters in bioprinting, which determine whether a printed construct can maintain its structure in a physiological environment [58,59].

Hydrogels for bioprinting of musculoskeletal tissues can be classified as natural or synthetic depending on their origin [60,61] and the most common ones used in the literature are summarized in Table 3.

Most natural hydrogels are based on components of the mammalian extracellular matrix (ECM), even though polymers from alternative sources, such as algae, are gaining interest [71]. Natural hydrogels

show significant bioactivity compared with synthetic materials due to the intrinsic presence of biomol-ecules used for signaling, adhesion, biocompatibility and self-remodeling [2]. While bioactive components are important for cell growth and differentiation, the application of natural materials can lead to batch-to-batch variability, immunogenic reactions and disease transmission [72]. Interestingly, it has been observed that natural polymers, such as hyaluronic acid, lam-inin, fibronectin and collagen, are more susceptible to cell-driven biodegradation [73].

It is important to note that as well as individual com-ponents of the ECM, decellularized ECM has similarly been successfully utilized as bioinks in the printing of tissue analogs [74,75]. Tissue decellularized ECM can be obtained using chemical, physical and biological treatments and provides an excellent representation of the natural ECM environment [76]. At the same time, decellularized ECM can lead to nonhomogeneous cell seeding and immune reactions, if cellular components are not fully removed [76]. Furthermore, decellulariza-tion treatments can damage the natural ECM and show poor mechanical properties in the material.

Synthetic materials are advantageous in terms of reproducibility and ease of processing compared with natural polymers. These polymers can sometimes lead to immunogenic reactions after partial degrada-tion, but this can be controlled and accelerated with the addition of matrix metalloproteinases, which show excellent biocompatibility [77,78]. Furthermore, syn-thetic materials can be tailored to form complexes with ECM proteins by covalent crosslinking.

These covalently bound molecules can be adhesion proteins or growth factors that enhance cell response within the hydrogels [58].

Challenges in bioprinting of musculoskeletal tissuesIn addition to the hardware used for bioprinting, mate-rial availability and their selection are significant chal-lenges and limitations for the success of bioprinting in musculoskeletal tissue fabrication. Materials/bioinks composed of naturally derived materials are limited in their application due to batch-to-batch variability and often lack the mechanical strength required to mimic the in vivo environment of native musculoskeletal tis-sues [59]. In the literature, this issue is often resolved by combining the natural inks with stronger biocompat-ible materials, such as PCL, PLA and PLGA [9,17], as described previously.

Furthermore, bioinks are presently limited by their printability and resolution. Materials are often required to be viscous in order to maintain the morphology of printed structures and improve mechanical strength,



but this can lead to blockages and unreliable material deposition. Importantly, this has adverse effects on the print quality and resolution, which can be detrimental to achieving the highly hierarchical structures in the tissues of the MSK.

Tuneable, synthetic bioinks can provide a wide range of desirable properties, including controlled mechanics, degradation and printability. However, the techniques required to synthesize and crosslink these materials can cause cytotoxicity and prevent the abil-ity to incorporate cells during the print process. For example, the monomers and photo initiators in some printable materials are toxic, but following UV cross-linking, the polymers formed are biocompatible and can support cell survival [79,80].

Further, printing of mechanically stable gela-tin in the form of GelMA traditionally requires UV light during the print process and can damage cells. Recently, this challenge has been overcome by devel-oping GelMA that can be crosslinked using visible light [81]. Crosslinking of natural bioinks can similarly be associated with cytotoxicity, but this can often be alleviated by using low crosslinker concentrations or replacing with nontoxic analogs. For example, calcium chloride is used for crosslinking alginate, whereas glu-taraldehyde is used for crosslinking collagen, gelatin or chitosan. Calcium chloride is not toxic at low con-centrations, while glutaraldehyde can be replaced with alternatives like genipin, during the crosslinking of collagen, gelatin and chitosan [82–84].

Hydrogels are the most utilized materials in bio-printing, due to their high water content and parallels to native ECM. However, these biomaterials show poor mechanical properties, so compromises must be made when considering characteristics such as composition, printability and mechanical strength. In the bioprint-ing of musculoskeletal tissues this has been overcome to an extent by incorporating multiple printing tech-nologies at the same time. This approach permits deposition of materials with good mechanical proper-ties and cell-laden bioinks, within a single e ngineered construct [21,26].

Furthermore, the fundamental layer-by-layer nature of most printing techniques leads to difficulties in producing complex and hollow structures. This can be resolved by incorporating sacrificial materials for structural support during the fabrication process but this also increases the technological complexity, cost and time of the printing. Once the resolution, repro-ducibility, speed and customization of current printing technologies have been defined and optimized, bio-printing can provide cost-effective and high-through-put systems for drug screening and tissue replacement. Importantly, robust methods for construct maturation

and long-term maintenance, as well as quality control measures for bioprinted tissues, need to be considered in parallel with the technological advances of printers. Regulatory concerns, such as the ethics of stem cells and the use of Class II medical devices are also key factors to contemplate, for the success of bioprinting approach. Furthermore, at present there are no specific 3D bioprinting regulations defined by the Medicines and Healthcare Products Regulatory Agency in the UK or US FDA in the USA.

Future perspectivePrinting of natural ECM-based materials embedded with cells provides significant advantages for tissue engineering, including biocompatibility, robust con-trol of cell distribution and density within the scaf-fold. Nevertheless, traditionally materials have been produced and seeded with cells afterward, and this can also be applied to printed constructs. In fact, this approach permits scaffolds with enhanced mechanical properties and resolution that natural materials often lack for musculoskeletal applications. Furthermore, co-printing of cell-laden and cell-free materials has been revealed to be beneficial for musculoskeletal tissue engineering applications [12,21,26]. All these approaches (cell-laden, cell-free and a combination of both) aim to obtain a complex, living tissue able to develop and mature, using the same fundamental technology.

The bioprinting literature suggests that the defini-tion of bioinks remains ambiguous. This is because various components such as cells, biomolecules, syn-thetic materials either alone or in combination, are all defined as bioinks by different groups. Consensus on bio-ink definition is required. With the development of new bioinks in the future, nanobiofabrication of organs will become a reality and this will help to alle-viate the increasing organ shortages worldwide. More specifically, technological advances in material sci-ence and engineering will permit versatility, nanoscale resolution and controllable distributing of cells and biomaterials, for a range of biomedical applications, i ncluding musculoskeletal repair.

The literature shows that bioprinting of muscles, tendons and ligaments is still a challenge; however, promising progress has been made in bioprinting of bone and cartilage. Stiff materials have been combined with natural cell-laden hydrogels to form composite constructs that are mechanically stable with the ability to mimic the native ECM environment of osteochon-dral tissues. Further, it has been demonstrated that these scaffolds can be combined with stem cells to per-mit osteochondral development in vivo [26]. Progress has even been made in the bioprinting of vascularized bone [34]. Biosensors for bone formation [85], and pro-



Executive summary

BackgroundSuccessful translation of tissue engineering and regenerative medicine research relies on efficient, robust and cost-effective fabrication techniques.Bioprinting is an automated additive manufacturing process that permits the fabrication of 3D structures by selectively depositing biological materials layer-by-layer.

Bioprinting & its role in musculoskeletal tissue fabricationBioprinting in the literature shows promise for musculoskeletal regeneration.Greater understanding of the native environment in tissues and organs is required to maintain cell viability.Printer hardware needs to be developed to combine cell health with desired biomaterials characteristics.

Musculoskeletal bioinks & their characteristicsBioinks can be printable biological materials that must be compatible with the biological, chemical and physical requirements of native tissues.To be suitable for musculoskeletal applications, a bioink must have suitable mechanical strength for cellular support and maintenance in 3D structures.A range of natural and synthetic materials have been used for musculoskeletal applications.Natural materials provide the biological properties for tissue development but lack the mechanical strength of bone and cartilage. Synthetic materials provide a solution to this challenge but are less bioactive.The most promising examples in the literature, combine cell-free and cell-laden printing.

Challenges in bioprinting of musculoskeletal tissuesBioinks cannot currently mimic the environments that cells experience in vivo.Printing living cells is challenging because there are multiple variables that need to be controlled and optimized when cells are combined with biomaterials.Hard tissues require mechanical strength that natural hydrogels often lack and synthetic materials can be detrimental to cell viability.

Future perspectiveBioinks must meet the following requirements: biocompatibility, biodegradability, good printability and sufficient mechanical strength for cellular support and maintenance of the 3D structures.The use of bioinks in bioprinting remains a developing and expanding multidisciplinary technology with substantial potential for the future successes of tissue engineering and regenerative medicine.

tein and DNA arrays of stem cells [86,87] have already been bioprinted, while next generation printable mate-rials for controlling osteochondral cellular microenvi-ronments are also in development. Nevertheless, these examples are all proof of concept studies and signifi-cant validation and development of next generation bioinks and their printing process is required in the future.

Fundamentally, the success of bioprinting in tis-sue engineering is heavily reliant on improvements in bioink properties, printing technologies, vasculariza-tion of tissues and controlled scaffold and cell matu-ration. Crucially for bone and cartilage applications, it is important to improve the mechanical proper-ties of bioinks and to maximize the resolution of the printed constructs. Innovative bioinks with benefits for musculoskeletal applications that are currently being developed include, dynamic switchable hydrogels with local variations in the density and size of collagen fibers throughout 3D tissues [88,89], and oxygen releasing b iomaterials [90,91].

ConclusionBioprinting promises to be an important tool to fab-ricate complex tissue and organs. However, there are significant challenges to be resolved in terms of techno-logical advances. Research to date has laid strong foun-dations and promise for the feasibility in manufactur-ing artificial organs, including musculoskeletal tissues. Bioprinting and the use of bioinks remain developing and expanding multidisciplinary fields of research with substantial potential for the future successes of tissue engineering and regenerative medicine.

Financial & competing interests disclosure



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