Bone Tissue Engineering Using 3D Printing in Nature Biotechnology

Materials Today � Volume 16, Number 12 �December 2013 RESEARCH

Bone tissue engineering using 3Dprinting

RESEARCH:Review

Susmita Bose*, Sahar Vahabzadeh and Am

W. M. Keck Biomedical Materials Research Lab, School of Mechanical and Materials Engineering, Washington State University, Pullman, WA 99164, USA

With the advent of additive manufacturing technologies in the mid 1980s, many applications benefited

from the faster processing of products without the need for specific tooling or dies. However, the

application of such techniques in the area of biomedical devices has been slow due to the stringent

performance criteria and concerns related to reproducibility and part quality, when new technologies

are in their infancy. However, the use of additive manufacturing technologies in bone tissue engineering

has been growing in recent years. Among the different technology options, three dimensional printing

(3DP) is becoming popular due to the ability to directly print porous scaffolds with designed shape,

controlled chemistry and interconnected porosity. Some of these inorganic scaffolds are biodegradable

and have proven ideal for bone tissue engineering, sometimes even with site specific growth factor/drug

delivery abilities. This review article focuses on recent advances in 3D printed bone tissue engineering

scaffolds along with current challenges and future directions.

IntroductionOsseous tissue, known as bone, is made of two different structures;

cancellous and cortical bone. Cancellous, or the inner part of bone,

is spongy in nature having 50–90 vol% porosity. However, cortical

bone is the dense outer layer of bone with less than 10 vol%

porosity. Both types of bone undergo dynamic remodeling, matura-

tion, differentiation, and resorption that are controlled via interac-

tions among osteocyte, osteoblast, and osteoclast cells [1].

Osteoblasts are primarily responsible for new bone formation while

osteoclasts are responsible for the resorption of old bone. Such a

dynamic process involving osteoclasts and osteoblasts is known as

bone remodeling, and is responsible for maintaining a healthy

bone. Bone is well known for its self-healing abilities [2]; however,

large-scale bone defects cannot be healed completely by the body

[3,4], and in most cases, external intervention is needed to restore

normal operations. Among different treatment options such as

autografts (bone taken from the same person’s body) and allografts

(bone tissue from a deceased donor), bone tissue engineering that is

focused on methods to synthesizeand/or regenerate bone to restore,

maintain or improve its functions in vivo [5,6] is becoming popular.

Successful application of bone tissue engineering can avoid

*Corresponding author:. Bose, S. ([email protected])

496 1369-7021/06 � 2013 Elsevier Ltd.Open a

challenges related to other treatment options involving different

materials such as autografts or allografts. Apart from material issues,

a clearunderstanding of biology involvingcells, extracellularmatrix

(ECM) and growth factors are pivotal in bone tissue engineering [7].

Scaffolds are an integral part of bone tissue engineering. Scaf-

folds are three dimensional (3D) biocompatible structures which

can mimic the ECM properties (such as mechanical support,

cellular activity and protein production through biochemical

and mechanical interactions), and provide a template for cell

attachment and stimulate bone tissue formation in vivo [3,5–7].

Besides chemistry, pore size, pore volume and mechanical strength

are critical parameters which define a scaffold’s performance. At an

early stage, bone ingrowth happens at the periphery of scaffolds

with a negative gradient in mineralization toward the inner parts

[4]. For continuous ingrowth of bone tissue, interconnected por-

osity is important. Open and interconnected pores allow nutrients

and molecules to transport to inner parts of a scaffold to facilitate

cell ingrowth, vascularization, as well as waste material removal

[4,6,8]. Since higher porosity increases surface area per unit

volume, the biodegradation kinetics of scaffolds can be influenced

by varying pore parameters. Biodegradation through a cell-

mediated process or chemical dissolution are both important to

ascertain stabilized repair and scaffold replacement with new bone

http://dx.doi.org/10.1016/j.mattod.2013.11.017ccess under CC BY-NC-ND license.

Materials Today � Volume 16, Number 12 �December 2013 RESEARCH

RESEARCH:Review

without any remnant [8]. A minimum pore size between 100 and

150 mm is needed for bone formation [4,9]; however, enhanced

bone formation and vascularization are reported for scaffolds with

pore size larger than 300 mm [9–11]. Pore size also plays an impor-

tant role in ECM production and organization. Poly(D,L-lactic acid)

(PDLLA) scaffolds with pore size 325 and 420 mm led to well-

organized collagen I network; whereas, smaller pore size of

275 mm prevented the human osteosarcoma-derived osteoblasts

to proliferate, differentiate and produce functional ECM [12]. Pore

volume also controls the permeability of nutrients to the scaffold

and their mechanical properties. Permeability in poly-e-caprolac-

tone (PCL) increased with higher pore volume and resulted in

better bone regeneration, blood vessel infiltration, and compres-

sive strength in vivo, when other pore parameters were kept the

same [13]. Apart from biological performance, the initial mechan-

ical properties and strength degradation rate should match that of

the host tissue for optimum bone healing [14]. The strength

degradation kinetics of porous scaffolds are highly affected by

pore size, geometry, and strut orientation with respect to the

loading direction [15,16]. Finally, surface properties such as chem-

istry, surface charge and topography also influence hydrophilicity

and in turn cell–material interactions for bone tissue ingrowth

[17–19].

Porous bone scaffolds can be made by a variety of methods.

Chemical/gas foaming [20], solvent casting, particle/salt leaching

[12,21], freeze drying [22], thermally induced phase separation

[23], and foam-gel [24] are some of those that have been used

extensively. However, pore size, shape, and its interconnectivity

cannot be fully controlled in these approaches. Moreover scaffolds

with tailored porosity for specific defects are difficult to manufac-

ture with most of these approaches [21–24]. Such scaffolds can be

designed and fabricated using additive manufacturing (AM)

approaches. Different AM approaches, for example, 3D printing

(3DP), solid freeform fabrication (SFF), rapid prototyping (RP), are

approaches that allow complex shapes for scaffolds’ fabrication

directly from a computer aided design (CAD) file [25–27]. The

concept of AM was first introduced by Chuck Hull in 1986 via a

process known as ‘stereolithography (SLA)’ [28,29].

Some of the commercially available AM techniques are 3DP

(ExOne, PA), fused deposition modeling (FDM, Stratasys, MN),

selective laser sintering (SLS, 3D Systems, CA), stereolithography

(3D Systems, CA), 3D plotting (Fraunhofer Institute for Materials

Research and Beam Technology, Germany), as well as various forms

of direct writing [27]. In all these AM approaches, 3D scaffolds are

created layer-by-layer without any part specific tooling or dies

[30,31]. These AM techniques can be classified as – (a) extrusion

(deformation + solidification), (b) polymerization, (c) laser-assisted

sintering, and (d) direct writing-based processes. Table 1 [32–59]

summarizes some of the AM techniques toward bone tissue engi-

neering applications including their advantages and disadvantages.

3D printing (3DP) – history and methodology3DP, a technology developed in the early 1990s at MIT (Cam-

bridge, MA) by Sachs et al. [60], is a powder-based freeform

fabrication method in which using a regular ink-jet print-head,

binders are printed on to loose powders in a powder bed. Early

research in this area was focused on rapid tooling using metals and

ceramics [61].

Fig. 1 shows a schematic representation of the 3DP process [62].

For bone tissue engineering, 3DP is useful for the direct fabrication

of scaffolds with tailored porosity from a CAD file. Before printing,

essential parameters such as powder packing density, powder

flowability, layer thickness, binder drop volume, binder saturation

and powder wettability need to be optimized to improve the

quality of the resultant part. Packing density is the relative density

of the powder bed after uniform spreading. To start a build,

enough powder should be packed homogeneously in a feed bed.

A set of rollers spread a layer of powder to a predetermined

thickness to create a powder bed. Powder flowability is critical

in this process as it determines the spreading ability. Flowability is

primarily determined by particle size, size distribution, surface

roughness and shape. The desired layer thickness is in part deter-

mined by geometry and powder characteristics. Thinner layers

cause binder penetration and excess spreading to other sites

resulting in poor resolution and tolerance. However, thick layers

need high saturation for the powders to bind [62].

The printhead sprays the binder across the build layer in several

passes, based on the instructions in the tool path file created

according to the CAD file. The binder, which can be organic or

water-based, locally binds the particles and hardens the wetted

area, or results in a reaction similar to the hydraulic setting

reaction in cements [63–65]. The binder drop volume and satura-

tion play crucial roles. The binder drop volume is the amount of

binder released from each nozzle per drop during printing, which

depends on the binder density and viscosity. By coordinating the

powder packing density and the drop volume, the binder satura-

tion data required for printing is obtained. For a constant packing

density, a higher drop volume demands a lower binder saturation

[62]. Low saturation can cause layer displacement during proces-

sing. The binder saturation also depends on the powder wettabil-

ity. The powder wettability, which is related to the powder particle

chemistry and surface energy, determines the printing accuracy

and the achievable tolerance [66]. While high wettability results in

extensive binder spreading, low wettability causes week powder-

binder integration [67].

After printing, the printed layer is moved under a strip heater to

allow the binder to dry out and prevent spreading between layers

[62]. This process is repeated until the printing of the designed part

is complete. Heat treatment is needed to complete the binder

reaction and increase the part green strength. Next step is depow-

dering, that is, the removal of loose powder from the printed body.

This is one of the major challenges for porous scaffolds in 3DP due

to the low green density of the part. Loose powder removal from

fine pores can easily crack a green part [68].

In general, a large variety of ceramic, metallic, polymeric, and

composite materials can be processed using 3DP; however, binder

selection and process parameter optimization are the keys to

successful part fabrication. In bone tissue engineering the advan-

tages of this method arise through the control of fine features

including interconnected porosity, no contamination issues

related to any second material for support structures and the direct

printing ability with both metallic and ceramic biomaterials

[65,69]. Fig. 2a shows some examples of 3D printed scaffolds with

different pore sizes. It is important to note that extensive optimi-

zation is needed to process good quality parts with 3DP for any

new material, a fundamental drawback for this approach.

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RESEARCH Materials Today � Volume 16, Number 12 �December 2013

TABLE 1

RP techniques for bone scaffold fabrication.

Technique Process details Processed materials for bonetissue engineering

Advantages (+) anddisadvantages (�)

Reference

3D Plotting/directink writing

! Strands of paste/viscous material(in solution form) extrusion based

on the predesigned structure

! Layer by layer deposition of

strands at constant rate,under specific pressure

! Disruption of strands according to

the tear of speed

! PCL! Hydroxyapatite (HA)

! Bioactive glasses

! Mesoporous bioactive

glass/alginate composite! Polylactic acid (PLA)/polyethylene

glycol (PEG)

! PLA/(PEG)/G5 glass

! Poly(hydroxymethylglycolide-co-e-caprolactone) (PHMGCL)

! Bioactive 6P53B glass

+:! Mild condition of process allows

drug and biomolecules (proteins

and living cells) plotting

�:! Heating/post-processing

needed for some materials restricts

the biomolecule incorporation

[32–38]

Laser-assisted

bioprinting

(LAB)

! Coating the desired material on

transparent quartz disk (ribbon)

! Deposition control by laser

pulse energy! Resolution control by distance

between ribbon/substrate, spot

size and stage movement

! HA

! Zirconia

! HA/MG63 osteoblast-like cell

! Nano HA! Human osteoprogenitor cell

! Human umbilical vein

endothelial cell

+:

! Ambient condition

! Applicable for organic,

inorganic materials and cells! Quantitatively controlled

! 3D stage movement

�:! Homogeneous ribbons needed

[39–42]

SLS ! Preparing the powder bed

! Layer by layer addition of powder! Sintering each layer according to

the CAD file, using laser source

! PCL

! Nano HA! Calcium phosphate (CaP)/poly

(hydroxybutyrate–co-

hydroxyvalerate) (PHBV)

! Carbonated hydroxyapatite(CHAp)/poly(L-lactic acid) (PLLA)

! PLLA

! b-Tricalcium phosphate (b-TCP)! PHBV

+:

! No need for support! No post processing is needed

�:! Feature resolution depends

on laser beam diameter

[43–48]

SLA ! Immersion of platform in a

photopolymer liquid! Exposure to focused light according

to desired design

! Polymer solidifying at focal point,non-exposed polymer remains liquid,

! Layer by layer fabrication by platform

moving downward

! Poly(propylene fumarate)

(PPF)/diethyl fumarate (DEF)! PPF/DEF-HA

! PDLLA/HA

! b-TCP

+:

! Complex internal features canbe obtained

! Growth factors, proteins and cell

patterning is possible�:! Only applicable for photopolymers

[49–52]

FDM ! Strands of heated polymer/ceramics

extrusion through nozzle

! Tricalcium phosphate (TCP)

! TCP/polypropylene (PP)

! Alumina (Al2O3)

! PCL! TCP/PCL

+:

No need for platform/support

�:! Material restriction due to need

for molten phase

[26,30,31,

53–58]

Robotic assisteddeposition/

robocasting

! Direct writing of liquid using anozzle

! Consolidation through liquid-to-

gel transition

! HA/PLA! HA/PCL

! 6P53B glass/PCL

+:! Independent 3D nozzle movement

! Precise control on thickness

! No need for platform/support

�:! Material restriction

[59]

RESEARCH:Review

3D printed bone scaffoldsTable 2 summarizes a few selected material-binder system combi-

nations for bone scaffolds using 3DP. Starch-based binders are one

of the candidates for bone replacement applications. These binders

are biocompatible and produce structures that have a mechanical

strength close to trabecular bone [71,72]. Structural designs and

post processing conditions both can influence the mechanical

properties of 3D-printed starch-based scaffolds [73]. 3D-printed

polyethylene (PE) scaffolds with 22.3–49.7% porosity have shown

498

a tensile strength up to 4 MPa, and no toxicity to human osteo-

blasts [74].

CaP ceramics are widely used in bone tissue engineering due to

their excellent bioactivity, osteoconductivity, and similarities in

composition to bone. Capillaries and vessel formation, and homo-

geneous osteoconduction from central channels, have previously

been observed in 3D-printed HA blocks [75]. The effect of pore size

on human fetal osteoblasts (hFOB) was studied with 3D-printed

TCP scaffolds [62]. The decrease in designed pore size from 1000 to

Materials Today � Volume 16, Number 12 �December 2013 RESEARCH

[(Figure_1)TD$FIG]

FIGURE 1

(a) 3D printing schematic using an inkjet printing system. (b) 3D printed CaP sintered structures fabricated at WSU.

TABLE 2

3D printed materials for bone tissue engineering.

Material Layer thickness Binder Reference

TCP 20 mm Aqueous based [62]

a/b-TCP modified with 5 wt% hydroxypropymethylcellulose 100 mm Water [64]

CaP mixture with Ca/P ratio of 1.7 100 mm 10% phosphoric acid [64]

Tetracalcium phosphate (TTCP), dicalcium phosphate and TCP 100 mm 25% citric acid [64]

HA 300 mm Schelofix (water soluble polymeric compound) [65,75]

TTCP/b-TCP 100 mm 25 wt% of citric acid [68]

TTCP/calcium sulfate dihydrate 100 mm 25 wt% of citric acid [68]

HA 100 mm No information [69]

TCP 100 mm No information [69]

Biphasic calcium phosphate (BCP) 100 mm No information [69]

a/b-TCP (final product: dicalcium phosphate dihydrate (DCPD)) No information 20% phosphoric acid [70]

Starch/PLLA + PCL No information Distilled water + blue dye [73]

High density PE (HDPE) 0.175 mm Maltodextrin + polyvinyl alcohol (PVA) [74]

SiO2–ZnO-doped TCP 20 mm Aqueous based [76]

PE or HDPE 0.175 mm Water based binder [77,78]

PLA No information Chloroform [79]

TCP (final product:DCPD) 0.1 mm 20% phosphoric acid [80]

HA/maltodextrin 0.175 Water based binder [81]

TTCP (final product: HA) 100 mm 0.5 mol/l Ca(H2PO4)2 + 10% H3PO4 [82]

TCP (final product: brushite) 100 mm 0.5 mol/l Ca(H2PO4)2 + 10% H3PO4 [8,82]

HA/A-W glass 0.1 mm Water based [83]

499

RESEARCH:Review

RESEARCH Materials Today � Volume 16, Number 12 �December 2013

[(Figure_2)TD$FIG]

FIGURE 2

(a) Photograph of the sintered 3D printed TCP scaffolds for mechanical strength and in vivo testing (small samples) [62]. (b) Compressive strength

comparison of the scaffolds sintered at 1250 8C in conventional and microwave furnaces (**p < 0.05, *p > 0.05, n = 10) [62]; (c) SEM micrographs of hFOB

cells showing the cell adhesion and proliferation on the scaffold surface and inside the 3D interconnected macro pores after 3 days of culture (white arrows

indicate cells): 500 mm (i) & (ii), and 750 mm (iii) & (iv) [62]; (d) SEM image of the pure TCP scaffold showing the surface morphology and designed macropore distribution [84]; (e) photomicrograph of 3DP pure (TCP) implants (i and iii), and Sr–Mg doped TCP implants (ii and iv) showing the new bone

formation inside the interconnected macro and intrinsic micro pores of the 3DP scaffolds after 4 and 8 weeks in rat distal femur model. Modified Masson

Goldner’s trichrome staining of transverse section. OB: old bone, NB: new bone and BM: bone marrow. Color description: Dark gray/black = scaffold; orange/

red = osteoid; green/bluish = new mineralized bone (NMB)/old bone [84]; and (f ) histomorphometric analysis of osteoid area fraction (osteoid area/total area,%) from 800 mm width and 800 mm height tissue sections (**p < 0.05, *p > 0.05, n = 8). Completely mineralized bone formation was observed in presence

of SrO and MgO in TCP after 12 weeks, hence no osteoid area was observed. All osteoid like bone was transformed into mineralized bone after 12 weeks in

doped TCP due to the presence of strontium and magnesium. Hence, there was no osteoid like bone left after 12 weeks in Sr–Mg-doped TCP [84].

RESEARCH:Review

750 and 500 mm resulted in an increase in proliferated cell density.

3D printed and microwave sintered b-TCP scaffolds fabricated

using 3DP are shown in Fig. 2a, showing interconnected macro

porosity across the sample. Fig. 2c (i–iv) presents the morphologies

of hFOB cells on scaffold surfaces and pore walls after 3 days of

culture showing good cell adherence and cell ingrowth into the

pores, suggesting that the scaffolds were non-toxic. A secondary

electron microscopy (SEM) image of the surface morphology and

the designed macro pore distribution in a pure TCP scaffold is

shown in Fig. 2d. New bone formation was observed at the

implant/host bone interface as well as inside the interconnected

macro and intrinsic micro pores after 4 and 8 weeks in both pure

and doped TCP as shown in Fig. 2e. However, more osteoid like

new bone formation was observed in SrO–MgO doped TCP scaffold

as shown in Fig. 2f. Histological evaluation and histomorpho-

metric analysis reveal that the treatment group (doped TCP scaf-

folds) facilitated higher osteoid like bone at an early stage, and

completely mineralized bone later, which could be essential for

fast bone healing and mineralization in vivo [62,84].

Further studies have shown that the addition of SiO2–ZnO

dopants to TCP scaffolds increases cell viability in different pore

size ranges [76]. The biocompatibility of 3D printed CaP ceramics

has also been studied using osteoclasts. Tartrate resistant acid

phosphatase (TRAP) staining, lacunae formation and microscopic

500

images confirmed the monocyte differentiation to multinuclear

osteoclast-like cells on a wide range of compositions [69]. It has

been shown that the use of phosphoric acid instead of polymeric

binders can improve both resolution and compressive strength

[64]. HA scaffolds with high surface areas showed no cytotoxicity

and adequate cell adhesion with MC3T3-E1 fibroblast cells [65]. In

addition to in vitro experiments, in vivo biocompatibility and

osteoconductivity of 3D-printed scaffolds showed that the 3D-

printed brushite and monetite cements with controlled open

porosity increased osteoconduction in vivo in a goat model [8].

3D-printed TCP samples with micro and macro-porosity also

facilitated osteogenesis in a rat femur model [53]. Cytotoxicity

results of MC3T3-E1 cells on two different bone cement based

compositions of TTCP/b-TCP and TTCP/calcium sulfate dihydrate

have been reported for bone tissue engineering. A wide range of

binders were used. It has been reported that the shortest hardening

time can be obtained between 20–40% of citric acid, and 30–40%

of lactic acid; however, a lower range of those binders and a

different concentration of sodium hydrogenphosphate with sul-

furic and phosphoric acids can be used to increase the hardening

time for the cements [68]. Fig. 3a and b show patient specific 3D

printed CaP implants. These results point to the application of 3DP

in a large variety of materials and structures for bone tissue

engineering scaffolds.

Materials Today � Volume 16, Number 12 �December 2013 RESEARCH

[(Figure_3)TD$FIG]

FIGURE 3

(a) 3D printed cranial segment [68], (b) general view of the implant bearing skull. Implants are fixed with miniplate osteosynthesis respectively bicortical

osteosynthesis (mandibular defect). The drill holes for screw insertion were made after the positioning of the implants using a common bone drill [70], (c)

representative macroscopic views of one half of bioceramic implant at retrieval, loaded with 56 ng copper [80].

RESEARCH:Review

Mechanical properties of 3D printed scaffoldsLow mechanical strength is a major challenge in porous scaffolds,

and is primarily controlled by pore volume. This is also true for 3D

printed ceramic scaffolds and limits their use only in non-load

bearing and low-load bearing applications. Optimized post proces-

sing approaches and compositional modifications can improve

mechanical properties of ceramic scaffolds. The compressive

strength of 3D printed TCP sintered scaffolds is shown in Fig. 2b.

Inagreement with observedshrinkage and increased density,micro-

wave sintering results in a higher compressive strength. The

strength of the scaffold increases with decreasing pore size or

volume, and a maximum strength of 10.95 � 1.28 MPa has been

observed for scaffolds with 500 mm pores, with 42% total open

porosity, when sintered at 1250 8C for 1 h in a microwave furnace

[62]. In another study, when a mixture of TTCP/b-TCP was sintered

at 1400 8C, it increased the strength of the 3D printed scaffold.

However, sintering a TTCP/calcium sulfate dihydrate composite

caused a decrease in the strength due to water release [68]. Tarafder

et al. reported an effective densification approach, using microwave

TABLE 3

Mechanical properties of 3D printed scaffolds.

Material Compressive

strength

(MPa)

TCP-sintered conventionally at 1250 8C 6.4

TCP-sintered using microwave at 1250 8C 10.9

TTCP/b-TCP 0.7

DLM infiltrated TTCP/b-TCP 76.1

Brushite

Monetite

Starch

PLLA/PCL infiltrated starch

TCP-sintered conventionally at 1250 8C 5.5

SiO2–ZnO doped TCP-sintered conventionally at 1250 8C 10.2

HA/A-W glass

HA/A-W glass-sintered at 1300 8C

HA

HA/bis-GMA

sintering compared to conventional heating, and improved the

mechanical properties of 3D-printed TCP scaffolds [62]. Bioactive

liquid phase sintering aids have also been reported to increase

strength. 3D printed HA/A-W glass, where the glassy phase is added

as a liquid phase sintering aid, showed an increase in strength from

1.27 MPa to 76.82 MPa when sintered at 1300 8C for 3 h [83]. The

enhancement of tensile properties was also found in PE scaffolds as a

result of thermally induced densification and binder degradation

[77]. To increase the strength of ceramic scaffolds without impairing

biological properties of scaffolds, another approach is monomer or

polymer infiltration. A mixture of bismethacrylated oligolactide

macromer (DLM-1), containing 10 wt% of 2-hydroxyethyl metha-

crylate has been used to increase the strength of scaffolds before and

after sintering [68]. The immersion of HA scaffolds in triethylene

glycol dimethacrylate (TEGDMA), 2,2-bis[4 (2-hydroxy-3thacryloy-

loxypropyloxy)-henyl] propane (bis-GMA) resulted in an increase of

the flexural strength byat least20 times [85]. Table 3 summarizes the

mechanical properties of 3D printed scaffolds tailored for bone

tissue engineering.

Compressive

stiffness

(MPa)

Compressive

yield strength

(MPa)

Bending

modulus

(GPa)

Bending

strength

(MPa)

Reference

[62]

[62]

[68]

[68]

5.2 [70]

3.9 [70]

11.15 1.12 [73]

55.19 1.77 [73]

[76]

[76]

0.35 1.27 [83]

34.1 76.82 [83]

0.4 0.69 [85]

6.2 50 [85]

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Bioprinting of tissue engineering scaffoldsApart from inorganic scaffold manufacturing, AM approaches are

also used to explore possibilities in fabricating scaffolds with live

cells and tissues. Organogenesis of liver tissue using 3D printed

PLLA/poly(lactic-co-glycolic acid) (PLGA) scaffolds has been inves-

tigated in vitro. It was shown that culturing a mixture of hepato-

cytes and endothelial cells on a channeled biodegradable scaffold

results in the desired tissue structure intrinsically [86]. In 3D fiber

deposition, a cell-laden viscous polymer paste was prepared and

printed using a syringe dispenser. Alginate hydrogel-embedded

multipotent stromal cells (MSCs)/chondrocytes were printed with

a high cell viability using this method. The incorporation of MSCs

and chondrocytes resulted in distinctive ECM formation both in

vitro and in vivo. In addition, an increase in strand distance was

shown to increase porosity with a lower elastic modulus. However,

changing the strand orientation from 908 to 458 increased the

elastic modulus [87]. Unique distribution and organization of

human umbilical vein endothelial cells (HUVECs) and mouse

embryonic fibroblast cells was obtained using gelatin methacrylate

scaffolds prepared by SLA [88,89]. 3D patterned human osteopro-

genitors (HOPs) and HA/HOPs fabricated by laser-assisted bio-

printing maintained osteoblastic phenotype and functionality

as evidenced by alkaline phasphatase (ALP) expression [39].

Growth factor and drug delivery using 3D printedscaffoldsThere are many growth factors such as vascular endothelial growth

factor (VEGF), fibroblast growth factors (FGFs) and bone morpho-

genic proteins (BMPs) that are important in bone tissue engineer-

ing. VEGF is an angiogenic protein which regulates endothelial

cell proliferation. FGFs are group of proteins, essential for the FGF

signaling pathway that induces angiogenesis through endothelial

and osteoblast cell proliferation, respectively [90]. BMPs, however,

induce osteogenesis through osteoprogenitors and mesenchymal

stem cell (MSC) differentiation to osteoblasts or binding to col-

lagen [62,91]. Bioprinting has become a versatile method in recent

years to create protein-based arrays. Bioprinting methods allow the

study of the effects of microenvironment changes due to the

aligned configuration to determine cell differentiation and align-

ment [92]. Muscle derived stem cells (MSDCs) cultured on printed

ECM containing BMP-2 indicated differentiation to osteogenic

lineage under myogenic conditions [91]. BMP-2 printed fibrin-

coated spun fibers regulated ALP as an osteoblast marker during

mouse myoblast cell line (C2C12) culture [91]. However, a delayed

BMP-2 administration from a 3D-printed HA block did not

enhance osteoinduction due to soft tissue ingrowth [75].

3D printed scaffolds have also been used for growth factor and

drug delivery to enhance bone growth in scaffolds. The localized

delivery of growth factors and drugs has attracted significant

attention due to the potential for dose reduction, controlled

release pattern, and the negligible side effects compared to sys-

temic delivery [93]. For scaffolds, pore size, connectivity and

geometry are effective parameters to control drug loading as well

as release rates in vivo [94]. Three different calcium phosphates

(CaPs) – brushite, monetite, and HA – were fabricated using 3DP as

shown in Table 2. Vancomycin hydrochloride, ofloxacin and

tetracycline hydrochloride were loaded onto these compositions

via immersion/vacuum impregnation. Drug absorption was

502

dependent on the specific surface area and the release followed

an exponential pattern. In addition, drug immersion in a poly-

lactide–polyglycolide (PLA/PGA) 50:50 polymer resulted in a

delayed release profile [82]. It was also shown that polymer incor-

poration in 3D-printed scaffolds could retard drug release kinetics

from first to zero order. In addition, vancomycin, heparin and

rhBMP-2 incorporation during printing revealed a reduction in

biological activity due to the degradation of drugs during spraying

through the nozzles [93]. Use of copper in DCPD scaffolds proved

that incorporation and release of copper can induce angiogenesis,

vasculogenesis and osteogenesis, as shown in Fig. 3c [80]. Fig. 4a

and b show a hexagonal gelatin methacrylate (GelMa) scaffold

seeded with HUVEC-green fluorescent protein (GFP), showing cell

spreading and organization [89]. Fig. 4c and d show vancomycin

release from 3D printed brushite and brushite/chitosan scaffolds

when a drug is loaded homogeneously, while Fig. 4e and f show

the release behavior from similar scaffolds when drug is loaded at

the center of scaffold [93]. Fig. 4g and h show various antibiotics

released from 3D printed calcium phosphate and brushite

matrices after immersion in a PBS buffer. The influence of PLA/

PGA polymer impregnation on vancomycin release from brushite

matrices is shown in Fig. 4i [82]. Overall, these studies show that

3D printed scaffolds can be used in drug delivery. There is a lot of

potential for direct printing of bone tissue engineering scaffolds,

but only if the reduction in biological activity of drugs/growth

factor can be minimized, and reproducibility and stability can be

assured. Today, bioprinting, and drug and growth factor delivery

using 3D printed tissue engineering scaffolds are still in their

infancies.

Future direction and challengesAM offers unique advantages toward part fabrication that are

needed for the production of small volumes or one of a kind

product manufacturing. Among the different AM techniques,

3DP is a versatile tool that has become popular for making scaf-

folds for bone tissue engineering. 3DP can fabricate scaffolds with

defined shapes, with controlled and interconnected porous struc-

tures. Although the process characteristics provide the opportu-

nity for the fabrication of almost all types of materials, the

selection of a suitable binder for 3DP is still a challenge, and

extensive optimization may be needed before high quality parts

can be made. Among different binders, organic binders work well,

however, they can affect the plastic parts of 3DP machines during

long term operation. The residue from binders may be difficult to

remove during sintering, an issue that may need special attention

for biomaterials. Moreover, to achieve the desired accuracy and

resolution in 3DP, a minimum distance between pores is necessary

which is dependent on powder characteristics and the build

parameters. The minimum distance requirement for a powder

based process makes it difficult to print highly porous scaffolds

with a sintered pore size below 300 mm [68].

Post processing is always required for 3DP processed parts.

Sintering or densification at high temperature is just one of them.

During sintering, parts shrink and the shrinkage is not necessarily

uniform throughout the part. Non-uniform shrinkage can cause

extensive cracking in parts and make them unusable. This is a

particular challenge for porous scaffolds. Since the outside part of

bone is a dense structure with�10% or less porosity while inside it

Materials Today � Volume 16, Number 12 �December 2013 RESEARCH

[(Figure_4)TD$FIG]

FIGURE 4

(a) Hexagonal gelatin methacrylate (GelMa) scaffold seeded with HUVEC-green fluorescent protein (GFP), (b) cells spreading and organization on scaffold

shown in (a) [89], (c) and (d) vancomycin release from 3D printed brushite and brushite/chitosan samples, when the drug is loaded homogeneously, (e) and(f ) vancomycin release from 3D printed brushite and brushite/chitosan samples, when the drug is loaded in the center of scaffold [93], (g) release of

vancomycin from different 3D printed calcium phosphate matrices after immersion in PBS buffer, (h) release of various antibiotics from brushite matrices

after immersion in PBS buffer, (i) influence of PLA/PGA polymer impregnation (10–50% polymer solution) on vancomycin release from brushite matrices [82].

RESEARCH:Review

is highly porous with>50% porosity, mimicking such structures is

very difficult using 3DP due to challenges related to non-uniform

shrinkage during sintering. Another post-processing challenge is

the removal of loose powders from interconnected pores inside the

part. This problem is magnified for parts with small pores, in

particular below 600 mm. Trapped powders inside the pores may

well sinter with the porous part making it less interconnected than

the designed part. Such problems with loose powders can reduce

the dimension of the pores after sintering.

Demand for processes such as 3DP will increase in the coming

years due to their ability to make custom medical devices that can

be tailored for patient specific and defect specific clinical needs.

Extensive process-property optimization is still needed to accom-

plish this goal. For ceramics, the most critical issue that needs

attention is the mechanical properties of porous scaffolds. Increas-

ing the porosity will decrease the strength of the scaffolds. Low

strength along with brittleness makes these scaffolds difficult to

even handle during processing. Resorbable polymer infiltration to

enhance strength and toughness in these scaffolds is one way to

minimize this problem; the use of resorbable glassy materials can

also help. Finally, printing live cells or adding growth factors/drugs

is another fascinating area of growth. However, most of the

challenges here are limited to survivability of the cells, viability

of the growth factors and drugs after printing. Although current

techniques let us build structures with similar composition to that

of tissue, we are still a long way from completely printing func-

tioning tissue [95]. More process-property optimization, in vitro

and in vivo research are needed in that direction to make any of

those approaches useful toward bone tissue engineering.

AcknowledgementsFinancial support from the US National Institute of Health under

the Grant Number (R 01 EB-007351) is acknowledged. Authors like

to also acknowledge Dr. Solaiman Tarafder for his help.

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