biomaterials 21 (2000) 2529}2543

Biomaterials 21 (2000) 2529}2543

Sca!olds in tissue engineering bone and cartilage

Dietmar W. HutmacherLaboratory for Biomedical Engineering, Institute of Engineering Science, Department of Orthopedic Surgery,

National University of Singapore, 10 Kent Ridge Crescent, Singapore 119260, Singapore

Abstract

Musculoskeletal tissue, bone and cartilage are under extensive investigation in tissue engineering research. A number ofbiodegradable and bioresorbable materials, as well as sca!old designs, have been experimentally and/or clinically studied. Ideally,a sca!old should have the following characteristics: (i) three-dimensional and highly porous with an interconnected pore network forcell growth and #ow transport of nutrients and metabolic waste; (ii) biocompatible and bioresorbable with a controllable degradationand resorption rate to match cell/tissue growth in vitro and/or in vivo; (iii) suitable surface chemistry for cell attachment, proliferation,and di!erentation and (iv) mechanical properties to match those of the tissues at the site of implantation. This paper reviews researchon the tissue engineering of bone and cartilage from the polymeric sca!old point of view. ( 2000 Elsevier Science Ltd. All rightsreserved.

Keywords: Tissue engineering of bone and cartilage; Design and fabrication of 3-D sca!old; Biodegradable and bioresorbable polymers

1. Introduction

Bone and cartilage generation by autogenous cell/tis-sue transplantation is one of the most promising tech-niques in orthopedic surgery and biomedical engineering[1]. Treatment concepts based on those techniqueswould eliminate problems of donor site scarcity, immunerejection and pathogen transfer [2]. Osteoblasts, chon-drocytes and mesenchymal stem cells obtained from thepatient's hard and soft tissues can be expanded in cultureand seeded onto a sca!old that will slowly degrade andresorb as the tissue structures grow in vitro and/orin vivo [3]. The sca!old or three-dimensional (3-D) con-struct provides the necessary support for cells to prolifer-ate and maintain their di!erentiated function, and itsarchitecture de"nes the ultimate shape of the new boneand cartilage. Several sca!old materials have been inves-tigated for tissue engineering bone and cartilage includ-ing hydroxyapatite (HA), poly(a-hydroxyesters), andnatural polymers such as collagen and chitin. Severalreviews have been published on the general propertiesand design features of biodegradable and bioresorbablepolymers and sca!olds [4}12]. The aim of this paper is tocomplete the information collected so far, with specialemphasis on the evaluation of the material and designcharacteristics which are of speci"c interest in tissueengineering the mesenchymal tissues bone and cartilage.

The currently applied sca!old fabrication technologies,with special emphasis on the so-called solid-free formfabrication technologies, will also be bench marked. Fi-nally, the paper discusses the author's research on thedesign and fabrication of 3-D sca!olds for tissue engi-neering an osteochondral transplant.

2. Polymer-based sca4old materials

The meaning and de"nition of the words biodegrad-able, bioerodable, bioresorbable and bioabsorbable(Table 1)*which are often used misleadingly in the tissueengineering literature*are of importance to discuss therationale, function as well as chemical and physical proper-ties of polymer-based sca!olds. In this paper, the polymerproperties are based on the de"nitions given by Vert [13].

The tissue engineering program for bone and cartilagein the author's multidisciplinary research curriculum hasbeen classi"ed into six phases (Table 2). Each tissueengineering phase must be understood in an integratedmanner across the research program*from the polymermaterial properties, to the sca!old micro- and macro-architecture, to the cell, to the tissue-engineered trans-plant, to the host tissue. Hence, the research objectives ineach phase are cross-disciplinary and the sub-projects arelinked horizontally as well as vertically.

0142-9612/00/$ - see front matter ( 2000 Elsevier Science Ltd. All rights reserved.PII: S 0 1 4 2 - 9 6 1 2 ( 0 0 ) 0 0 1 2 1 - 6

Table 1De"nitions given by Vert

Biodegradable are solid polymeric materials and devices which breakdown due to macromolecular degradation with dispersion in vivo butno proof for the elimination from the body (this de"nition excludesenvironmental, fungi or bacterial degradation). Biodegradable poly-meric systems or devices can be attacked by biological elements so thatthe integrity of the system, and in some cases but not necessarily, of themacromolecules themselves, is a!ected and gives fragments or otherdegradation by-products. Such fragments can move away from theirsite of action but not necessarily from the body.

Bioresorbable are solid polymeric materials and devices which showbulk degradation and further resorb in vivo; i.e. polymers which areeliminated through natural pathways either because of simple "ltrationof degradation by-products or after their metabolization. Bioresorptionis thus a concept which re#ects total elimination of the initial foreignmaterial and of bulk degradation by-products (low molecular weightcompounds) with no residual side e!ects. The use of the word &bio-resorbable' assumes that elimination is shown conclusively.

Bioerodible are solid polymeric materials or devices, which show sur-face degradation and further, resorb in vivo. Bioerosion is thus a con-cept, too, which re#ects total elimination of the initial foreign materialand of surface degradation by-products (low molecular weight com-pounds) with no residual side e!ects.

Bioabsorbable are solid polymeric materials or devices, which candissolve in body #uids without any polymer chain cleavage or molecu-lar mass decrease. For example, it is the case of slow dissolution ofwater-soluble implants in body #uids. A bioabsorbable polymer can bebioresorbable if the dispersed macromolecules are excreted.

Table 2The research program for tissue engineering bone and cartilage classi-"ed into six phases

I*Fabrication of bioresorbable sca!oldII*Seeding of the osteoblasts/chondrocytes populations into the

polymeric sca!old in a static culture (petri dish)III*Growth of premature tissue in a dynamic environment

(spinner #ask)IV*Growth of mature tissue in a physiologic environment (bioreactor)V*Surgical transplantationVI*Tissue-engineered transplant assimilation/remodeling

The "rst stage of tissue engineering bone or cartilagebegins with the design and fabrication of a porous 3-Dsca!old, the main topic of this review paper. In general,the sca!old should be fabricated from a highly biocom-patible material which does not have the potential toelicit an immunological or clinically detectable primaryor secondary foreign body reaction [9]. Furthermore,a polymer sca!old material has to be chosen that willdegrade and resorb at a controlled rate at the same timeas the speci"c tissue cells seeded into the 3-D constructattach, spread and increase in quantity (number ofcells/per void volume) as well as in quality. Currently, thedesign and fabrication of sca!olds in tissue engineeringresearch is driven by three material categories: I. Regula-tory approved biodegradable and bioresorbable poly-mers (Table 3), such as collagen, polyglycolide (PGA),

polylactides (PLLA, PDLA), polycaprolactone (PCL),etc. II. A number of non-approved polymers, such aspolyorthoester (POE), polyanhydrides, etc. which arealso under investigation. III. The synthesis of entrepre-neurial polymeric biomaterials, such as poly (lactic acid-co-lysine), etc., which can selectively shepherd speci"c cellphenotypes and guide the di!erentiation and prolifer-ation into the targeted functional premature and/or ma-ture tissue.

In general, polymers of the poly(a-hydroxy acids)group undergo bulk degradation. The molecular weightof the polymer commences to decrease on day one (PGA,PDLA) or after a few weeks (PLLA) upon placement inan aqueous media [12]. However, the mass loss does notstart until the molecular chains are reduced to a sizewhich allows them to freely di!use out of the polymermatrix [14]. This phenomenon described and analyzed indetail by a number of research groups [15}18], results inaccelerated degradation and resorption kinetics until thephysical integrity of polymer matrix is compromised. Themass loss is accompanied by a release gradient of acidicby-products.

In vivo, massive release of acidic degradation andresorption by-products results in in#ammatory reac-tions, as reported in the bioresorbable device literature[19}22]. If the capacity of the surrounding tissue toeliminate the by-products is low, due to the poor vas-cularization or low metabolic activity, the chemical com-position of the by-products may lead to local temporarydisturbances. One example of this is the increase of osmo-tic pressure or pH manifested through local #uidaccumulation or transient sinus formation from "berreinforced polyglycolide pins applied in orthopedic sur-gery [21]. Potential problems of biocompatibility intissue engineering bone and cartilage, by applying degra-dable, erodable, and resorbable polymer sca!olds, mayalso be related to biodegradability and bioresorbability.Therefore, it is important that the 3-D sca!old/cell con-struct is exposed at all times to su$cient quantities ofneutral culture media, especially during the period wherethe mass loss of the polymer matrix occurs.

The incorporation of a tricalciumphosphate (TCP)[23], hydroxyapatite (HA) [24] and basic salts [15] intoa polymer matrix produces a hybrid/composite material.These inorganic "llers allow to tailor the desired degra-dation and resorption kinetics of the polymer matrix.A composite material would also improve biocompatibil-ity and hard tissue integration in a way that ceramicparticles, which are embedded into the polymer matrix,allow for increased initial #ash spread of serum proteinscompared to the more hydrophobic polymer surface [9].In addition, the basic resorption products of HA or TCPwould bu!er the acidic resorption by-products of thealiphatic polyester and may thereby help to avoid theformation of an unfavorable environment for the cellsdue to a decreased pH [15,23,24].

2530 D.W. Hutmacher / Biomaterials 21 (2000) 2529}2543

Tab

le3

Pro

per

ties

ofbi

ores

orb

able

and

bioer

oda

ble

poly

mer

s

Poly

mer

Com

pariso

nof

mec

han

ical

proper

ties

ofbio

erodab

lean

dbio

reso

rbab

lepol

ymer

s

Deg

radat

ion

and

reso

rption

pro

cess

via

hyd

roly

sis

Mol

ecula

rw

eigh

tlo

ss/loss

of

mec

han

ical

proper

ties

(inm

ont

h)!

Mas

slo

ss(in

mont

h)!

Ref

eren

ces

(sca!old

s)R

efer

ence

s(m

edic

alde

vice

)A

rea

of

appl

icat

ion

Pro

duc

tsw

ith

regu

lato

ryap

pro

val

Poly

(L-lac

tide

)#

##

Bulk

erosion

9}15

36}48

46,4

8,49

,60}64

,70}

7219

,20,

24O

rthop

edic

Sur

gery

,Ora

lan

dM

axill

ofac

ial

Sur

gery

Fix

Sorb

Sys

tem

(scr

ews,

nai

ls,pin

s)N

eo"x

(scr

ews,

nails,

pins)

Poly

(L-lac

tide

-co-D

,L-lac

tide)

70/3

0#

#Bulk

erosion

5}6

12}18

22,2

3O

ralan

dM

axill

ofac

ial

Surg

ery,

Ort

hop

edic

Sur

gery

Res

orP

in,L

ead"x

Mac

roSor

bSys

tem

(scr

ews

and

pla

tes,

mes

h,nai

ls,pi

ns)

Poly

Pin

Poly

(L-lac

tide

-co-

glyc

olid

e)10

/90

##

Bulk

erosion

1}2

3}4

28,6

3Sut

ure

Per

iodo

nta

lSur

gery

,Surg

ery,

Vic

rylSutu

re,

Vic

rylM

esh

Poly

glyc

olid

e#

##

Bulk

erosion

0.5}

13}

46,

30}35

,41,

65O

rthop

edic

Sur

gery

Bio"x

Poly

(D,L

-lac

tide

)#

Bulk

erosion

1}2

5}6

6,31

,34,

49,6

0Poly

(D,L

-lac

tide

-co

-gly

colid

e)85

/15

#Bulk

erosion

1}2

4}5

53}56

,60,

70

Poly

(D,L

-lac

tide

-co

-gly

colid

e)75

/25

#Bulk

erosion

1}2

4}5

49,6

1

Poly

(D,L

-lac

tide

-co

-gly

colid

e)50

/50

##

Bulk

erosion

1}2

3}4

28

Poly

capr

ola

cton

e#

Bulk

and

surfac

eer

osion

9}12

24}36

29D

rug

deliv

ery

Cap

rano

r

Poly

orth

oes

ter

##

Surfac

eer

osion

4}6

12}18

Poly

anhyd

rides

##

Surfac

eer

osion

4}6

12}18

59A

nim

alex

perim

ents

!Mole

cula

rw

eigh

tan

dm

ass

loss

vary

dep

endin

gon

fact

ors

such

asch

emic

alst

ruct

ure

and

com

position

;pre

senc

eofi

onic

grou

psan

dof

side

group

defe

cts;

con"gu

ration

oft

he

stru

cture

mole

cula

rw

eigh

tan

dm

ole

cula

rw

eigh

tdistr

ibution

(pol

ydisper

sity

);pr

esen

ceof

low

mole

cula

rw

eigh

tco

mpo

nen

ts(m

ono

mer

s,olig

omer

s,so

lven

ts,so

ften

ers,

dru

gs,gr

ow

thfa

ctors

,et

c.);

pro

duc

tion

and

man

ufac

turing

pro

cedu

res

and

thei

rpr

oces

spar

amet

ers,

impl

ant

des

ign,

ster

iliza

tion

met

hod

,m

orp

holo

gy(a

mor

phous

vers

us

sem

i-cr

ysta

llin

e,pre

sence

ofm

icro

stru

cture

san

dst

ress

withi

nth

eco

mpo

nen

ts),

tem

pering,

stor

age,

impla

nt

site

.#

##

,good

;#

#,av

erag

e;#

,poor

.

D.W. Hutmacher / Biomaterials 21 (2000) 2529}2543 2531

Fig. 1. Graphical illustration of the complex interdependence of molecular weight loss and mass loss of a strategy I 3D sca!old matrix plotted againstthe time frame for tissue engineering a cartilage/bone transplant. (A) Fabrication of bioresorbable sca!old; (B) seeding of the osteoblast/cartilagepopulations into the polymeric sca!old in a static culture (petri dish); (C) growth of premature tissue in a dynamic environment (spinner #ask);(D) growth of mature tissue in a physiologic environment (bioreactor); (E) surgical transplantation; (F) tissue-engineered transplant assimila-tion/remodeling.

Control of the hydrodynamic and biochemical envi-ronment is essential for the successful in vitro engineeringof 3-D sca!old/tissue constructs for potential clinical use[25]. Computer-controlled bioreactors that continuouslysupply physiological nutrients and gases, serve to regu-late the required cell/tissue culture conditions for a longperiod of time. After the in vitro culturing of the 3-Dsca!old/tissue construct, the degree of remodeling andcell/tissue replacement of the bone/cartilage transplantby the host tissue has to been taken into consideration[26]. Cell and tissue remodeling is important for achiev-ing stable mechanical conditions and vascularization atthe host site. Hence, the 3-D sca!old/tissue constructshould maintain su$cient structural integrity during thein vitro and/or in vivo growth and remodeling process.The degree of remodeling depends on the host anatomyand physiology [26]. The polymer selection from a ma-terial science point of view is based on two di!erentstrategies in regard to the overall function of the sca!old.

2.1. Strategy I

In the "rst strategy (Fig. 1), the physical sca!old struc-ture supports the polymer/cell/tissue construct from thetime of cell seeding up to the point where the hard tissuetransplant is remodeled by the host tissue. In the case ofload-bearing tissue such as articular cartilage and bone,

the sca!old matrix must serve an additional function; itmust provide su$cient temporary mechanical support towithstand in vivo stresses and loading. In Strategy I re-search programs, the material must be selected and/ordesigned with a degradation and resorption rate suchthat the strength of the sca!old is retained until the tissueengineered transplant is fully remodeled by the hosttissue and can assume its structural role.

Bone is able to remodel in vivo under so-called physio-logical loading [27]. It is a requirement that the degrada-tion and resorption kinetics have to be controlled in sucha way that the bioresorbable sca!old retains its physicalproperties for at least 6 months (4 months for cell cultur-ing and 2 months in situ). Thereafter, it will start losing itsmechanical properties and should be metabolized by thebody without a foreign body reaction after 12}18 months(Fig. 1). The mechanical properties of the bioresorbable3-D sca!old/tissue construct at the time of implantationshould match that of the host tissue as closely as possible[7]. It should posses su$cient strength and sti!ness tofunction for a period until in vivo tissue ingrowth hasreplaced the slowly vanishing sca!old matrix.

Thompson et al. [28] studied a poly(D,L-lactide-co-glycolide) matrix under cyclic compressive loading.They concluded that changes in surface deformation andmorphology suggest that the compressive loading ini-tially collapses and sti!ens the polymer matrix. The de-


Fig. 2. Graphical illustration of the complex interdependence of molecular weight loss and mass loss of a strategy II 3D sca!old matrix plotted againstthe time frame for tissue engineering a cartilage/bone transplant. (A) Fabrication of bioresorbable sca!old; (B) seeding of the osteoblast/cartilagepopulations into the polymeric sca!old in a static culture (petri dish); (C) growth of premature tissue in a dynamic environment (spinner #ask);(D) growth of mature tissue in a physiologic environment (bioreactor); (E) surgical transplantation; (F) tissue-engineered transplant assimila-tion/remodeling.

crease in molecular weight is slowed down due to thereduction of surface area from hydrolysis, until thematrix architecture no longer accommodates the mech-anical loading and begins to lose its integrity. Conclus-ively, in Strategy I the sca!old architecture has towithstand mechanical loading in vitro and in vivo.

2.2. Strategy II

For Strategy II (Fig. 2), the intrinsic mechanical prop-erties of the sca!old architecture templates the cell prolif-eration and di!erentiation only up to the phases wherethe premature bone or cartilage is placed in a bio-reactor. The degradation and resorption kinetics of thesca!old are designed to allow the seeded cells to prolifer-ate and secrete their own extracellular matrix in the staticand dynamic cell seeding phase (weeks 1}12), while thepolymer sca!old gradually vanishes leaving su$cientspace for new cell and tissue growth. The physical sup-port by the 3-D sca!old is maintained until the engineer-ed tissue has su$cient mechanical integrity to supportitself.

Di!erent research groups [29}33] have shown ina number of studies that a nonwoven mesh made ofpolyglycolide "bers o!ers degradation and resorptionkinetics for Strategy II. However, the challenge for thegrown cell/tissue construct is to have similar mechanicalproperties to the host bone and cartilage. Ma and Langer[34] showed that cartilage which was cultured for

7 month in a bioreactor reached 40% of the mechanicalproperties of natural cartilage. The next phase of thoseresearch programs will be to evaluate how these tissue-engineered cartilage transplants assimilate and remodelin vivo [35]. In an in vivo model, one of the majorproblems from a biomechanical and clinical view point, isthe primary mechanical stabilization of cartilage trans-plants [28]. This aspect will be discussed below, undersca!old design, in more detail.

3. Sca4old design

Skeletal tissue, such as bone and cartilage, is usuallyorganized into 3-D structures in the body [36]. For therepair and generation of hard and ductile tissue, such asbone, sca!olds need to have a high elastic modulus inorder to be retained in the space they were designated for;and also provide the tissue with adequate space forgrowth [37]. If the 3-D sca!old is used as a temporaryload-bearing device (Strategy II), the mechanical proper-ties would maintain that load for the required time with-out showing symptoms of fatigue or failure. Therefore,one of the basic problems from a sca!old design point ofview, is that to achieve signi"cant strength the sca!oldmaterial must have su$ciently high interatomic and in-termolecular bonding, but must have at the same timea physical and chemical structure which allows for hy-drolytic attack and breakdown.


Fig. 3. The Cologne cathedral was built in the 18th and 19th century bygothic architects who used a stone skeleton structure of stone columnsand ribs supported by arches and buttresses.

Ingber and his group [38,39] design their sca!oldsbased on a concept which they named tensegrity.Geodesical 3-D constructs were designed by applyingtensegrity so that the entire sca!old structure evenlydistributes and balances mechanical stresses. The walls,layers or struts that make up the interconnecting sca!oldframework are connected into triangles, pentagons orhexagons, each of which can bear tension or compres-sion. However, the mechanical rational of the tensegritydesign concept has been known for centuries in the areaof civil engineering. Gothic architects used a stone skel-eton structure of stone columns and ribs supported byarches and buttresses to build cathedrals (Fig. 3).

Another point, which has to be focused on is thedi!usion of nutrients into the 3-D sca!old. Although, aninterconnected macropore-structure of 300}500 lm en-hances the di!usion rates to and from the center ofa sca!old, transportation of the nutrients and by-prod-ucts is not su$cient for large sca!old volumes. A #uid-dynamic microenvironment provided by a bioreactor canmimic the interstitial #uid conditions present in naturalbone and cartilage in a macroporous sca!old architec-ture. Bioreactors permit in vitro culture of larger andbetter organized 3-D cell communities than can beachieved using standard tissue culture techniques [40].

For tissue engineering a bone transplant, the creationof a vascularized bed ensures the survival and function ofseeded cells, which have access to the vascular system fornutrition, gas exchange, and elimination of by-products[41]. The vascularization of a sca!old may be compro-mised by purely relying on capillary ingrowth into theinterconnecting pore network from the host tissue.In situ, the distance between blood vessels and mesen-chymal cells are not larger then 100 lm [42]. Therefore,the time frame has to be taken into account for thecapillary system to distribute through larger sca!oldvolume. It may also be possible to control the degree andrate of vascularization by incorporating angiogenic andanti-angiogenic factors in the degrading matrix of thesca!old.

From a biomechanical and clinical point of view, thetissue-engineered bone or cartilage transplant shouldallow for a mechanically secure and stable "xation on orto the host tissue [29]. For bone, the currently availablemedical devices, such as pins, screws, and plates might beused. However, the integration of a device-like part intothe 3-D sca!old design can be advantageous. The ration-ale for such an innovative design concept will be for thetissue engineering of an osteochondral bone transplant isdescribed below.

4. Sca4old fabrication

A number of fabrication technologies have beenapplied to process biodegradable and bioresorbablematerials into 3-D polymeric sca!olds of high porosityand surface area. The conventional techniques for scaf-fold fabrication include "ber bonding, solvent casting,particulate leaching, membrane lamination and meltmolding (Table 4). Several papers have reviewed the pastand current research on sca!old fabrication techniques[43}45]. However, none of those papers has directlycompared the 3-D sca!old-processing technologies forthe tissue engineering community. From a sca!old designand function view point each processing methodologyhas its pro and cons. It is the aim of this paper toaggregate the compiled information and to present thisdata in a comprehensive form. Table 4 summarizes thekey characteristics and parameters of the techniques cur-rently used. The aim of this part of the review paper is toassist research teams with their choice for a speci"c 3-Dsca!old-processing technology by providing the informa-tion needed to determin the critical parameters. As dis-cussed above, at present the challenge in tissueengineering bone and cartilage is not only to design, butalso to fabricate reproducible bioresorbable 3-D scaf-folds, which are able to function for a certain period oftime under load-bearing conditions.

Solvent casting, in combination with particle leaching,works only for thin membranes or 3-D specimens with


Tab

le4

Cur

rently

appl

ied

3Dsc

a!old

fabrica

tion

tech

nolo

gies

Fab

rica

tion

tech

nolo

gyPro

cess

ing

Mat

eria

lpro

pert

ies

required

for

pro

cess

ing

Sca!

old

des

ign

and

repro

duc

ibility

Ach

ieva

ble

por

esize

inlm

Poro

sity

in%

Arc

hitec

ture

Ref

eren

ce

Solv

entca

stin

gin

com

binat

ion

with

part

icula

rle

achin

g

Cas

ting

Sol

ubl

eU

ser,

mat

eria

lan

dte

chniq

ue

sens

itiv

e30}30

020}50

Spher

ical

por

es,sa

ltpar

ticl

esre

mai

nin

mat

rix

47

Mem

bra

ne

lam

inat

ion

Sol

ventbo

ndin

gSol

ubl

eU

ser,

mat

eria

lan

dte

chniq

ue

sens

itiv

e30}30

0(

85Ir

regu

lar

por

est

ruct

ure

48

Fab

rica

tion

ofnon

-wove

nC

ardin

g,N

eedlin

g,Pla

tepre

ssin

gFib

res

Mac

hine

contr

olle

d20}10

0(

95In

su$

cien

tm

echan

ical

prop

erties

30}35

,41

,63}65

Mel

tm

ould

ing

Mou

ldin

gTher

mop

last

icM

achi

ne

contr

olle

d50}50

0(

80Ext

rusion

inco

mbin

atio

nw

ith

part

icul

arle

achi

ng

Ext

rusion

thro

ugh

die

sTher

mop

last

icM

achi

ne

contr

olle

d(

100

(84

Spher

ical

por

es,sa

ltpar

ticl

esre

mai

nin

mat

rix

49

Em

ulsio

nfree

zedry

ing

Cas

ting

Sol

ubl

eU

ser,

mat

eria

lan

dte

chniq

ue

sens

itiv

e(

200

(97

Hig

hvo

lum

eofin

ter-

conn

ecte

dm

icro

por

est

ruct

ure

54}56

Ther

mal

lyin

duce

dph

ase

sepa

ration

Cas

ting

Sol

ubl

eU

ser,

mat

eria

lan

dte

chniq

ue

sens

itiv

e(

200

(97

Hig

hvo

lum

eofin

ter-

conn

ecte

dm

icro

por

est

ruct

ure

58}61

Sup

ercr

itic

al-#

uid

tech

nol

ogy

Cas

ting

Am

orp

hous

Mat

eria

lan

dte

chniq

ue

sens

itiv

e(

100

10}30

Hig

hvo

lum

eofnon

inte

rconne

cted

mic

ropor

est

ruct

ure

53

Sup

ercr

itic

al-#

uid

tech

nol

ogy

inco

mbi

nat

ion

with

part

icle

leac

hin

g

Cas

ting

Am

orp

hous

Mat

eria

lan

dte

chniq

ue

sens

itiv

eM

icro

pore

s(50

mac

ropor

es(

400

(97

Low

volu

me

ofno

n-int

erco

nnec

ted

mic

ropor

est

ruct

ure

com

bin

edw

ith

inte

rconn

ecte

dm

acro

por

est

ruct

ure

54

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Fig. 5. Schematic illustration of the fused deposition modeling (FDM)process.

Fig. 4. Polymer disks with a diameter of 500 lm and 40 lm thicknessallow to stack a 3D sca!old with a porosity of 98%.

very thin wall sections: otherwise, it is not possible toremove the soluble particles from within the polymermatrix [46]. Mikos et al. [47], using the above-describedtechnology, fabricated porous sheets and laminated themto 3-D structures. Chloroform was used on the attach-ment interface for the lamination process. This fabrica-tion technology is time consuming because only thinmembranes can be used. Another disadvantage is that thelayering of porous sheets allows only a limited number ofinterconnected pore networks. Solvent-casted polymer-salt composites have also been extruded into a tubulargeometry [48]. The disadvantages of the above technolo-gies include the extensive use of highly toxic solvents,time required for solvent evaporation (days-to-weeks),the labor intensive fabrication process, the limitation tothin structures, residual particles in the polymer matrix,irregularly shaped pores, and insu$cient interconnectivity.

The supercritical #uid-gassing process has been knownfor many years in the non-medical polymer industry [49]as well as in the pharmaceutical community [50]. Thistechnology is used to produce foams and other highlyporous products. The polymers which can be used forthis technology have to have an high amorphous frac-tion. The polymer granules are plasticized due to theemployment of a gas, such as nitrogen or carbon dioxide,at high pressures. The di!usion and dissolution of the gasinto the polymer matrix results in a reduction of theviscosity, which allows the processing of the amorphousbioresorbable polyesters in a temperature range of30}403C [51]. The supercritical #uid-gassing technologyallows the incorporation of heat sensitive pharmaceut-icals and biological agents. However, on average only10}30% of the pores are interconnected [51,52]. Harriset al. [53] combined this technology with particulateleaching to gain a highly interconnected void network.The researchers could control porosity and pore size byvarying the particle/polymer ratio and particle size.

Whang et al. [54,55] developed a protocol for thefabrication of aliphatic polyester-based sca!olds by usingthe emulsion freeze-drying method. Sca!olds with poros-ity greater than 90%, median pore sizes ranging from 15to 35 lm with larger pores greater than 200 lm werefabricated. The sca!old pore architecture was highly in-terconnected which is necessary for tissue ingrowth andregeneration [54]. Based on their results from an animalexperiment, the interdisciplinary group proposed a scaf-fold design concept which results in in vivo bone regen-eration based on hematoma stabilization [56]. Theauthors compare their in vivo bone engineering conceptto the induction phase of fracture healing. The osteop-rogenitor cells which are in the blood of the osseouswound are embedded in the sca!old microarchitecturevia the hematoma. The multipotent cells di!erentiate toosteoblasts due to the presence of growth factors whichare released by the host bone. However, the emulsionfreeze-drying method is user and technique sensitive. The

fabrication of a truly interconnecting pore structure de-pends on the processing method and parameters as wellas on the used equipment [57,58].

Several groups [57}60] studied thermally inducedphase separation technology to process polymeric 3-Dsca!olds. This technique has been used previously tofabricate synthetic membranes for non-medical applica-tions. The method has been extensively applied in the"eld of drug delivery to fabricate microspheres, whichallows the incorporation of pharmaceutical and biolo-gical agents, such as bone morphogenetic proteins(BMPs) into the polymer matrix. In general, the micro-and macro-structure is controlled by varying the polymermaterial, polymer concentration, quenching temper-ature, and solvents. However, current research showsthat the method, similar then emulsion freeze-dryingtechnique, is user and technique sensitive and that theprocessing parameters have to be well controlled. Namand Park [57] as well as Zhang and Ma [58] fabricatedpolymer and polymer/HA specimens with a porosity of


Fig. 6. (a) 3D sca!old systems of various porosity and pore geometry fabricated by FDM. Magni"cation ]7.5, scale bar represents 1mm. (i)}(iii)lay-down pattern: 0/903; nozzle tip: 0.016A; porosity: 50, 68, 75%; (iv)}(vi) 0/903; 0.010A; 50, 68, 75%; (vii)}(viii) 0/60/1203; 0.016A; 68, 75%; (ix) 0/60/1203;0.010A; 80%; (x)}(xii) 0/60/1203; 0.010A; 50, 68, 75%. (b) Left: cross-sectional view of freeze-fractured PCL sca!old with lay-down pattern0/72/144/36/1083. Right: plan view of same specimen. (c) Left: top view of PCL-HA sca!old with lay-down pattern 0/60/1203. The materialcomposition consists of 80% PCL and 20% HA by weight. Right: a close-up view of the same specimen at a higher magni"cation, shows that HAparticles are at the sca!old surface. (d) Freeze-fracture cross-sectional surface (]23) of a bioerodable sca!old designed for a bone/cartilage interface.A 0/60/1203 and a 0/903 lay-down pattern of the roads shown in the upper and lower portion respectively of the SEM picture. Despite being fabricatedwith di!erent lay-down patterns, the porosity (67%) of both portions could be made identical. However, by applying FDM the porosity of eachportion of the 3D sca!old can be varied according to the road spacing for individual layers.

up to 95%. At present, only pore sizes of up to 100 lmcan be reproducibly fabricated by thermally inducedphase-separation technology.

A number of textile technologies have the potential todesign and fabricate highly porous sca!olds [61]. Yet,only so-called non-woven mesh-like designs have been

used to tissue engineer bone and cartilage [62}64]. Excel-lent results in tissue engineering cartilage have beenachieved by using non-woven meshes composed of poly-mer "bers of PGA, PGA/PDLA, and PGA/PLLA. Thiswork has been reviewed by Freed et al. [65] and will notbe discussed here. In general, non-woven constructs can


Fig. 6. (Continued).

be only used for Strategy II since their physical proper-ties do not allow load-bearing applications.

All the above-described technologies except the mem-brane-lamination method, do not allow the fabrication ofa 3-D sca!old with a varying multiple layer design. Sucha matrix architecture is advantageous in instances wheretissue engineers want to grow a bi- or multiple tissueinterface, e.g. an articular cartilage/bone transplant.

Rapid prototyping technologies as well as so-called &wa-fer stacking systems' [66] (Fig. 4) have the potential todesign a 3-D construct in a multi-layer design within thesame gross architectural structure.

In engineering literature [67] Rapid prototyping Tech-nologies (RP) also called Solid Free Form fabrication(SFF) methods are de"ned as a set of manufacturingprocesses that are capable of producing complex-free


Fig. 7. Schematic illustration of the surgical placement of a tissueengineered a bone/cartilage interphase. The dental cylinder implant-like design of the bony part of the sca!old allows to apply a press-"tbetween the host bone and the cylindrical portion of the sca!old.

form parts directly from a computer-aided design(CAD) model of an object without part speci"c tooling or knowledge. Unlike machining processes such asmilling, drilling, etc., which are subtractive in nature, RPsystems join together liquid, powder and sheet materialsto form parts. Layer by layer, RP machines fabricateplastic, wood, ceramic and metal objects using thin hori-zontal cross sections directly from a computer-generatedmodel.

Rapid prototyping technologies, such as 3-D printing(3-DP) and fused deposition modeling (FDM) allow thedevelopment of manufacturing processes to create por-ous sca!olds that mimic the microstructure of livingtissue. Three-dimensional printing*developed at theMassachusetts Institute of Technology [68]*is alsoa rapid prototyping technology which has been used toprocess bioresorbable sca!olds for tissue engineering ap-plications [69,70]. The technology is based on the print-ing of a binder through a print head nozzle ontoa powder bed, with no tooling required. The part is builtsequentially in layers. The binder is delivered to thepowder bed producing the "rst layer, the bed is thenlowered to a "xed distance, powder is deposited andspread evenly across the bed, and a second layer is built.This is repeated until the entire part, e.g. a porous scaf-fold, is fabricated. Following treatment, the object isretrieved from the powder bed and excess unbound pow-der is removed. The speed, #ow rate and even dropposition can be computer controlled to produce complex3-D objects [63]. This printing technique permits CAD)and custom-made fabrication of bioresorbable hybridsca!old systems. The entire process is performed underroom-temperature conditions. Hence, this technologyhas great potential in tissue engineering applications.Biological agents, such as cells, growth factors, etc., canbe incorporated into a porous sca!old without inactiva-tion if non-toxic binders, e.g. water can be used [71].Unfortunately, aliphatic polyesters can be only dissolvedin highly toxic solvents, such as chloroform, methylenechloride, etc. To date, only bioresorbable sca!olds with-out biological agents within the polymer matrix and incombination with particle leaching have been processedby 3-D printing. In addition, the mechanical propertiesand accuracy of the specimen manufactured by 3-Dprinting have to be signi"cantly improved [72].

The FDM process forms 3-D objects from a CAD "leas well as digital data produced by an imaging sourcesuch as computer tomography (CT) or magnetic reson-ance imaging (MRI) [73]. The process begins with thedesign of a conceptual geometric model on a CAD work-station. The design is imported into a software, whichmathematically slices the conceptual model into horizon-tal layers. Toolpaths are generated before the data isdownloaded to the FDM hardware. The FDM extrusionhead (Fig. 5) operates in the X and > axes while theplatform lowers in the Z-axis for each new layer to form.

In e!ect, the process draws the designed model (sca!old)one layer at a time.

Thermoplastic polymer material, 1.78 mm in diameter,feeds into the temperature-controlled FDM extrusionhead where it is heated to a semi-liquid state. The headextrudes and deposits the material in ultra-thin layersonto a "xture less base. The head directs the materialprecisely into place. The material solidi"es, laminating to


Fig. 8. (a) Schematic illustration of a perfusion culture chamber for tissue engineering a bone/cartilage interphase; (b) schematic sketch of thebioreactor concept for the tissue engineering of bone and cartilage simultaneously, in the one device like sca!old architecture.

the preceding layer. Parts are fabricated in layers, wherea layer is built by extruding a small bead of material, orroad, in a particular lay-down pattern, such that the layeris covered with the adjacent roads. After a layer is com-pleted, the height of the extrusion head is increased andthe subsequent layers are built to construct the part. Inthe past, non-medical and medical FDM users couldonly use a few non-resorbable polymeric materials, suchas polyamide, ABS, and other resins. The broadening of

the research and application scope of FDM results ina need/demand to evaluate the critical factors for usingthe new polymeric and composite materials on FDMsystems [74].

At present, the author's multidisciplinary group hasbeen able to evaluate the parameters to process a numberof potential sca!old materials, such as PCL andPCL/HA by FDM. To our knowledge, we are the "rstgroup to report the processing of bioresorbable sca!olds


(Fig. 6a}d) for tissue engineering applications usingFDM.

5. Tissue engineering of a articular cartilage}boneinterface

Articular cartilage repair, regeneration and generationhave been reviewed by Buckwalter and Mankin [75] aswell as Newman [76]. Both the reports have concludedthat in the last two decades, clinical and basic scienti"cinvestigations have shown that the implantation ofarti"cial matrices, growth factors, perichondrium, andperiosteum, can stimulate the formation of cartilaginoustissue in osteochondral and chondral defects in synovial#uids. However, the available evidence indicates that theresults vary considerably among the individuals, and thatthe tissues formed using these treatment regimes do notduplicate the composition, structure, and mechanicalproperties of normal articular cartilage [77}81]. In addi-tion, none of those matrix designs could be securely "xedunder load-bearing conditions to the osteochondral bonewhich is a conditio sine qua non from a biomechanicaland clinical point of view.

The author's multidisciplinary group has concep-tualized a rationale for tissue engineering a load-bearingosteochondral transplant (Fig. 7). The sca!old design,material and fabrication technology, as well as thebioreactor design (Fig. 8a and b), enables the tissueengineering of bone and cartilage simultaneously, in theone sca!old architecture. The device-like design of thebony sca!old part allows a secure bone-to-bone "xationof the articular bone}cartilage interface. The concept hasbeen discussed in detail elsewhere [82].

6. Conclusions

The application of regulatory approved biomaterialsto design and fabricate 3-D sca!olds has stronglysupported the drive for the establishment of tissueengineering research. Reviewing the experimental andclinical studies, it can be concluded that the idealsca!old and matrix material for tissue engineeringbone and cartilage has not yet been developed. Ingeneral, the tissue engineers do not implement inno-vative sca!old "xation features into their design concept.However, from a clinical point of view, the secureand user friendly transplant "xation is a conditio sinequa non. The shape of a mesenchymal tissue such asa bone and articular cartilage is often critical to itsfunction. Ideally, a polymeric sca!old material shouldpermit application of a solid-free form fabrication tech-nology, so that a porous sca!old with any desired 3-Dgeometry can be designed and fabricated by using CTand MRI data.

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