Biomedical production of implants by additive electro-chemical and ...

uceorenced asnts

anditro

ilityy, awasha-

s totionWG theing,riesical

CIRP Annals - Manufacturing Technology xxx (2012) xxx–xxx

ible

ctor

ntly

rial-

cess

wth

sed

een

IRP.

G Model

CIRP-916; No. of Pages 21

Biomedical production of implants by additive electro-chemical and physicalprocesses

Paulo Bartolo (2)a,*, Jean-Pierre Kruth (1)b, Jorge Silva c, Gideon Levy (1)a,d, Ajay Malshe (2)e,Kamlakar Rajurkar (1)f, Mamoru Mitsuishi (1)g, Joaquim Ciurana h, Ming Leu (1)i

a Centre for Rapid and Sustainable Product Development, Polytechnic Institute of Leiria (IPL), Leiria, Portugalb Division PMA (Production Engineering), University of Leuven (KU Leuven), Leuven, Belgiumc Centro de Tecnologia de Informacao Renato Archer, Campinas, Brazild Inspire AG, Institute for Rapid Product Development, St. Gallen, Switzerlande Department of Mechanical Engineering, University of Arkansas, Fayetteville, AR, USAf Department of Mechanical Engineering, University of Nebraska, Lincoln, NE, USAg Department of Mechanical Engineering, The University of Tokyo, Tokyo, Japanh Department of Mechanical Engineering and Industrial Construction, University of Girona, Girona, Spaini Department of Mechanical Engineering and Aerospace Engineering, Missouri University of Science and Technology, Rolla, MO, USA

1. Introduction

The ageing population, high expectations for a better quality oflife and the changing lifestyle of modern society require improved,more efficient and affordable health care. This poses newchallenging problems regarding the increasing number of implantsrequired, new diseases to be treated (e.g., Parkinson’s andAlzheimer’s) and organ shortage problems. On the other hand,some medical devices ideally should survive without experiencingany failures for the patent’s lifetime.

The loss or failure of an organ or tissue is a frequent and costlyproblem in health care. Today, treatments include either trans-planting organs from one individual to another or performingsurgical reconstructions by transferring tissue from one location inthe patient’s body into the diseased site. The disparity between theneed and availability of donor tissues has motivated thedevelopment of tissue engineering approaches aimed at creatingcell-based substitutes of native tissues [16,17,151].

biocompatible materials, cells and growth factors to prodbiological structures for tissue engineering applications’’ [22]. Mrecently, in a meeting sponsored by the American National ScieFoundation in the spring of 2008, biomanufacturing was define‘‘the design, fabrication, assembly and measurement of bio-elemeinto structures, devices, and systems, and their interfacing

integration into/with larger scale structures in vivo or in v

environment such that heterogeneity, scalability and sustainabare possible.’’ In 2009, during the 59th CIRP General AssemblCollaborative Working Group (CWG) on biomanufacturing

established based on three main pillars: Biofabrication, Biomectronics and Biodesign, and Assembly. The goal of this CWG icontribute to a coherent strategy for the development, disseminaand exploitation of biomanufacturing. To pursue this goal, the Caims to optimise current technologies and develop new ones inareas of computer-integrated surgical systems, tissue engineerbio-informatics and nano diagnosis/medicine, based on the theoand the technologies established in each CIRP Scientific Techn

A R T I C L E I N F O

Keywords:

Additive manufacturing

Biomedical

Health care

A B S T R A C T

Biomanufacturing integrates life science and engineering fundamentals to produce biocompat

products enhancing the quality of life. The state-of-the-art of this rapidly evolving manufacturing se

is presented and discussed, in particular the additive electrical, chemical and physical processes curre

being applied to produce synthetic and biological parts. This fabrication strategy is strongly mate

dependent, so the main classes of biomaterials are detailed. It is explained the potential to pro

composite materials combining synthetic and biological materials, such as cells, proteins and gro

factors, as well the interdependences between materials and processes. The techniques commonly u

to increase the bioactivity of clinical implants and improve the interface characteristics betw

biological tissues and implants are also presented.

� 2012 C

Contents lists available at SciVerse ScienceDirect

CIRP Annals - Manufacturing Technology

journal homepage: http: / /ees.elsevier.com/cirp/default .asp

cusxity

of

To address some of these demanding issues, a new scientificdomain called biomanufacturing emerged in 2005, during theBiomanufacturing Workshop hosted by Tsinghua University in Chinaand defined as ‘‘the use of additive technologies, biodegradable and

reefacenly

ring* Corresponding author. Tel.: +351 244 569 441; fax: +351 244 569 444.

E-mail address: [email protected] (P. Bartolo).

Please cite this article in press as: Bartolo P, et al. Biomedical pprocesses. CIRP Annals - Manufacturing Technology (2012), http:/

0007-8506/$ – see front matter � 2012 CIRP.

http://dx.doi.org/10.1016/j.cirp.2012.05.005

Committee (STC).This review follows the establishment of the CIRP CWG’s fo

on current healing and repairing strategies. Despite the compleassociated with the design, fabrication and implantationappropriate medical implants, this paper addresses only thcritical topics: biomaterials, manufacturing processes and surtreatments for the fabrication of clinical implants, the obiomedical implants considered here (Fig. 1). Biomanufactu

roduction of implants by additive electro-chemical and physical/dx.doi.org/10.1016/j.cirp.2012.05.005


mailto:[email protected]


http://www.sciencedirect.com/science/journal/00078506


is a

dureprodthosphythe

applthe

(perdedipolyconsthesand

righapprand

releand

impusedbetwtionFinathe

betw

2. M

Msurfto restrumissmon

Ito t‘‘perplacbodreplbodpermimpespespinbiodScafto fa

P. Bartolo et al. / CIRP Annals - Manufacturing Technology xxx (2012) xxx–xxx2

G Model


Plepr

strongly material- and process-dependent fabrication proce- in which materials not commonly used in conventionaluction engineering are considered. The main characteristics ofe materials strongly determine the electro-chemical andsical additive manufacturing processes to be used, as well asapplication range of these production technologies. Theication context of this work is detailed in Section 2, wheremain characteristics of the considered clinical implants

manent and temporary) are described. Section 3 is fullycated to the four main classes of biomaterials (metals,mers, ceramics and composites) used to produce theidered implants. Understanding of the main properties ofe biomaterials and the interdependences between materialsbiological tissues is fundamental not only for selecting the

t material for a specific application but also for selecting theopriate manufacturing process. Materials such as hydrogelbiomaterials/cells composites are also introduced due to theirvance. Section 4 introduces the most relevant electro-chemicalphysical additive processes used for the production of clinicallants. The main characteristics, applications and materials

by each of these technologies are explained. The integrationeen materials (Section 3), processes (Section 4) and applica-

s (Section 2) are summarised at the end of Section 4 (Table 7).lly, in Section 5, some techniques are introduced to enhancebioactivity and the establishment of strong connectionseen biological tissues and implants.

edical implants

edical implants are devices placed either inside or on theace of the body to accomplish some particular function, such asplace, assist or enhance the functionality of some biological

cture(s). Many implants are prosthetics, intended to replaceing body parts, while other implants deliver medication,itor body functions, or provide support to organs and tissues.

mplants are classified as permanent or temporary. Accordinghe United States Food and Drug Administration (FDA), amanently implantable device is a device that is intended to be

2.1. Biodegradable implants

Degradable implants or scaffolds serves as temporary skeletonsto accommodate and stimulate new tissue growth (Fig. 2). Theyplay a major role in tissue engineering representing the initialbiomechanical support for cell attachment, differentiation andproliferation [16,17,142,148].

An ideal scaffold must satisfy the following requirements[16,17,92,142,144,182]:

� Biocompatibility. Both raw and processed materials shouldinteract positively with the host environment without elicitingadverse host tissue responses.� Biodegradability. Scaffolds must degrade into non-toxic products

with a controlled degradation rate that matches the regenerationrate of the native tissue. The in vivo degradation process ofpolymeric scaffolds is influenced by different and oftenconflicting variables, such as those related to the material’sstructure (i.e., chemical composition, molecular weight andmolecular weight distribution, crystallinity, morphology, etc.), itsmacroscopic features (i.e., implant shape or size, porous shape,size and interconnectivity, etc.) and environmental conditions(i.e., temperature, pH of the medium, presence of enzymes orcells and tissues).

The chemical degradation of polymers may principally proceedvia either degradation by biological agents (enzymes), hydrolyticdegradation (hydrolysis), which is mediated by water, or acombination of both coming into contact with living tissue.Several authors have investigated the degradation process of awide range of biomaterials [55,68,80,197]. Lee et al. [129], Sunget al. [197], Agrawal et al. [2,3] and Lu et al. [146] studied the

Fig. 1. Classification of biomedical implants.

Fig. 2. Tissue engineering process, involving seed cells on scaffold, culturing in vitro

and implant into the patient.

Adapted from [142].

ed into a surgically or naturally formed cavity of the humany for more than one year to continuously assist, restore, orace the function of an organ system or structure of the humany throughout the useful life of the device.’’ Examples of

anent implants include stents and hip implants. Temporarylants are commonly used in sports and medical surgeries,cially in shoulder and knee ligamentous reconstruction andal reconstructive surgery [203]. They are usually made ofegradable polymers like screws, suture threads and plates.folds are permanent or temporary porous structures implantedvour tissue or bone regeneration.

ase cite this article in press as: Bartolo P, et al. Biomedical pocesses. CIRP Annals - Manufacturing Technology (2012), http:/

degradation of poly(lactic-co-glycolic acid) (PLGA) and polyca-prolactone (PCL) and found that the degradation rate depends onthe molecular weight and hydrophobicity. Lam et al. [124] showedthat the hydrolytic degradation of PCL scaffolds is governed bytheir high molecular weights, crystallinity, hydrophobicity, sur-face-to-volume and porosity. On the other hand, incorporatingcertain other materials, such as calcium phosphate, significantlyincreases the degradation rate [124]. Domingos et al. [55,56]observed the in vitro degradation of PCL scaffolds in simulatedbody fluid (SBF) and the phosphate buffer solution (PBS) for 6months. Results show a more significant degradation process of the



areuseicalfectrialvedand44]

rate thentlys offies

over

as it

P. Bartolo et al. / CIRP Annals - Manufacturing Technology xxx (2012) xxx–xxx 3

G Model


scaffolds in SBF than in PBS, due to the formation of calciumphosphate deposits on PCL scaffolds, which decreases thedegradation kinetics.

Many environmental factors are involved in the gradualdegradation of calcium phosphate ceramic after implantation,including physicochemical processes (dissolution-precipitation)and the effects of various cell types [83,135,213]. Cells such asfibroblasts and osteoblasts degrade ceramics by either phagocy-totic mechanisms or an acidic mechanism with a proton pump toreduce the pH of the microenvironment, resorbing these syntheticsubstrates (osteoclasts) [83,141]. The cellular mechanisms ofcalcium phosphate degradation are modulated by several para-meters, such as the properties of the ceramic itself, theimplantation sites and the presence of various proteins (cytokines,hormones, vitamins, ions, etc.) [83,141].

� Appropriate porosity, pore size and pore shape. Generally, a highlevel of porosity is required (>90%) because it increases thesurface area, enabling high cell seeding efficiency, migration andproliferation, as well as neovascularisation [121]. Pore size playsan important role in terms of cell adhesion/migration, vascular-isation and new tissue ingrowth [17,85,99,128,136,164,180].Macro-pores (i.e. >50 mm) are of an appropriate scale toinfluence tissue function, while micro-pores (i.e. <50 mm)influence cell function (e.g. cell attachment), given thatmammalian cells typically are 10–20 mm in size and nano-porosity refers to pore architectures or surface textures at thenano-scale level (i.e. 1–1000 nm) [93]. Optimum macro-poresizes of 20 mm have been reported for fibroblast ingrowth, 20–125 mm for hepatocytes and 100–250 mm for the regeneration ofbone. If the macro-pores are too large or too small, cells will failto spread and form networks throughout the scaffold. Smallerpores enhance cell adhesion and differentiation in vitro, whilebigger pores promote higher cell adhesion and viability in vivo. Itis important to define an optimal macro-pore size range forsupporting cell and tissue ingrowth. These limits vary greatlydepending upon cell type and the culture conditions, but ingeneral they fall in the range of 100–500 mm. Pore intercon-nectivity (100% interconnected network of internal channels arerequired) is also a critical parameter in terms of cell viability andtissue regeneration, maximising the diffusion and exchange ofnutrients and the eliminations of waste. Pore interconnectivitycan be measured by either determining the flow rate of fluidflowing through the scaffold or using techniques, such asmercury intrusion porosimetry, micro-computed tomography(mCT) or image analysis [103,204].� Bioactive. Scaffolds should be bioactive, promoting and guiding

cell proliferation, differentiation and tissue growth. This can beachieved by adding growth factors and functionalizing thescaffold with proteins or adhesion-specific peptide sequences,which often resemble the extracellular matrix providing appro-priate signals to cells.� Mechanical strength. Scaffolds are required to withstand both in

vitro manipulation and stresses in the host tissue environment. In

vitro, engineered culture tissue constructs should maintain theirmechanical properties to preserve the required space for cellgrowth and matrix formation. For in vivo applications, it isimportant that scaffolds mimic as closely as possible themechanical properties of the native tissue in order to provide

processes [4,64]. However, a few suitable techniques

available to sterilise biodegradable polyester scaffolds becathey are susceptible to degradation and/or morphologdegeneration under high temperatures and pressures. The efof sterilisation on both surface topography and mateproperties should be considered. It was previously obserthat ETO and, for example g-radiation, induces degradation

shrinking of poly(a-hydroxyesters) [27,86]. Cottam et al. [observed that gamma irradiation significantly decreased the

of degradation of PCL discs, although the rates depended oninitial mass of the polymer. Gamma irradiation also significaincreased the mechanical yield stress, but not the failure stresPCL. Andrews et al. showed that sterilisation method modithe topography of scaffolds (Fig. 4) [5].

Fig. 3. Schematic representation of the mechanical contribution of a scaffold

time as it degrades, and the mechanical contribution of the new host tissue

forms in the presence of appropriate mechanical loading [11].

Fig. 4. AFM images of electrostatically spun polyurethane scaffolds: (A) virgin; (B)

UV-ozone sterilised; (C) ETO sterilised [5].

a temporary mechanical support for tissue regeneration. Initially,the scaffold must withstand all stresses and loads in the hosttissue environment before gradually transferring them to theregenerated tissue (Fig. 3).� Adequate surface finish. This guarantees a good biomechanical

coupling between the scaffold and the tissue.� Easily manufactured and sterilised. Implants should be rapidly

produced with high accuracy and repeatability and easilysterilised by exposure to high temperatures, UV light, g-radiation, plasma, ethylene oxide (ETO) gas, or by immersionin a sterilisation agent, remaining unaffected by either of these

Please cite this article in press as: Bartolo P, et al. Biomedical production of implants by additive electro-chemical and physicalprocesses. CIRP Annals - Manufacturing Technology (2012), http://dx.doi.org/10.1016/j.cirp.2012.05.005


2.2.

Tpracuse

restmus

ADevof oeithcan

The

tiontimebiocon t

Mhighdeveresesubjman

IapplaccojoinvalvimpCosmaestprosbucoInfoBioCfunc

Fprosthe

segmits fto th

Imecpossterianessgreaingr

Iexce

Fig. 5Feder

Fig. 6coop


G Model


Plepr

Permanent implants

he first evidence of head surgery dates back to 8000 BC. Thetice of immobilising members began around 3000 BC, and theof metal and other material implants, such as bone, for dentaloration and cranioplasty can be observed in archaeologicaleums all over the world.ccording to the National Institutes of Health Consensus

elopment Conference, an implant is ‘‘a medical device madene or more biomaterials intentionally placed within the body,er totally or partially buried beneath an epithelial surface’’ thatbe in contact with tissue for a significant period of time [12].biocompatibility of implants is time dependent and applica-

related: a material can be biocompatible for a short period of but incapable of long-term contact with tissue. It can also beompatible for one application but not for another, dependinghe surrounding tissue.

edical implants have been used for about 50 years; today, a percentage of people have such a device, especially inloped countries where access to health care is affordable and

arch/technology more developed. However, all implants areect to failure, and long-term data concerning their perfor-ce and host response are vital for improvements.

mplants can be functional or cosmetic. A broad variety ofications exist for functional implants, which are intended tomplish the organ’s function. Examples include hip and other

t implants, vascular prostheses, artificial ligaments, heartes, cages, spinal fusion and spacer devices, some cochlearlants, artificial hearts and dental implants [10,33,112,206].

etic implants can be used to provide shape or improvehetical functions for breast, nose, ear, hand, and foottheses, among many others. Fig. 5 shows a cosmetic, ormaxillofacial prosthesei, developed by the Renato Archer

rmation Technology Center (CTI, Brazil) using the concept ofAD [102]. Beyond aesthetic purposes, this prosthesis showedtional results, improving patient speech and feeding.ig. 6 shows the sequence used for producing of a cranial hollowthesis made of polymethylmethacrylate–PMMA. The design ofprosthesis was designed using a Magnetic Resonanceented model of the lesion area (duramater membrane), and

unction was to keep intracranial pressure stable, giving shapee external area.

mplants can be mechanically fixed with screws, pins, wires,hanical interference, glue or cement. Biological fixation is alsoible with surface treatments and improvements with bioma-ls and chemical modifications. Augmenting of surface rough-

and area with microspheres, meshes, and filaments is also oft importance when biological attachment through tissueowing is expected.mplants are expected to be durable and stable with anllent interface when in contact with tissue. For example, a

biomaterial in contact with blood, as in stents and heart valves,cannot permit any blood disorder, such as attachment or coagulumformation. On the other hand, metal stems for hip implants withbiological attachment are expected to have a better tissue growingresponse to promote fixation. Implants must be designed to be asleast invasive as possible during installation and maintenance.

Some problems related to implant failure, such as loosing,breaking and migration, are usually caused by the wrong design(not designed for a specific patient), material selection, application,manipulation, installation and misuse. Some of the causes arebiologically related, such as infection and excessive tissueresponses, while others are mechanically-related, such as stressconcentration, micro-displacements, surface scratches affectingthe passivation layer (surface film created to reduce chemicalreactivity), and wear generating particles or debris. Fatigue andcorrosion are very common in metal implants, and wear is frequentin polymeric parts of implants; therefore, to successfully accom-plish its task, an implant has to be designed considering severalcriteria, such as physicochemical properties, strength, fatigueendurance, wear resistance and material dimensional stability,always taking into account the application and surrounding tissue.

3. Materials

Biomaterials are materials that interface with biological entities[29,174,202,214]. The National Institutes of Health ConsensusDevelopment Conference defined a biomaterial as ‘‘any substance(other than a drug) or combination of substances, synthetic ornatural in origin, which can be used for any period of time, as awhole or as a part of a system which treats, augments, or replacesany tissue, organ, or function of the body’’ [214]. A distinctivedifference between a biomaterial over other materials is its benigncoexistence with a biological system with which it interfaces [81].The use of bioinert, bioactive materials (first generation ofbiomaterials) is an important response to the growing medicalneeds of a rapidly ageing population. Subsequently, the biomater-ials field began to shift from bioinert materials to bioactivematerials, which can elicit controlled actions and reactions withinthe body. Currently, four classes of biomaterials are used:

� Acellular tissue matrices (biological scaffolds);� Metallic materials;� Ceramic material;� Polymers (naturally derived and synthetic polymers);

Biological scaffold materials, composed of an extracellularmatrix (ECM), are not considered in this paper, as they are notprocessed through the manufacturing processes considered here.The preparation of these biological matrices commonly involves acombination of physical (freezing, direct pressure, sonication andagitation), chemical and enzymatic methods to remove cell bodiesfrom the remaining extra-cellular matrix [45,67].

3.1. Metals

Metals and their alloys, due to their mechanical reliability,strength, stiffness, toughness and impact resistance, were used forload-bearing implants, such as hip and knee prostheses andfracture fixation wires, pins, screws, and plates (Fig. 7). Metals have

. Lip, nose and part of maxilla prostheses developed at CTI in cooperation with

al University of Minas Gerais (Brazil).

. Cranial reconstruction of the front-temporo-parietal area developed at CTI in

eration with the Sobrapar Hospital (Brazil). Fig. 7. (A) Hip implant, (B) knee implant.

ase cite this article in press as: Bartolo P, et al. Biomedical production of implants by additive electro-chemical and physicalocesses. CIRP Annals - Manufacturing Technology (2012), http://dx.doi.org/10.1016/j.cirp.2012.05.005


oneand

alsoure

ng’s84],

andties,

sivem-

gO),os-ics

d toiumted.ce-

heirons,ical

icsatesuef a

The ofichith

ium. 9

and


G Model


also been used as parts of artificial heart valves, vascular stents,and pacemaker leads [112].

Important characteristics to be considered for medical applica-tions are biocompatibility, appropriate mechanical properties,corrosion resistance and structural integrity [175]. Metallicbiomaterials are classified as inert materials because they elicitminimal tissue response. In physiological environments, metalscan suffer from corrosion, thus releasing ions, which may reducebiocompatibility and put at risk the use of implants.

Major metals used in medical applications include commer-cially pure titanium and its alloys (a+b alloys, Ti–6Al–4V, Ti–Al–Nb and b-Ti alloys), cobalt-based alloys (Co–Cr–Mo, Co–Ni–Cr–Mo,Co–Cr–W–Ni), stainless steel (primarily type 316L), Ni–Ti alloys,Au-based materials, and Ag–Sn alloys [31,41,150].

Titanium has one of the highest strength-to-weight ratios andcorrosion resistance of metals [86]. It has excellent biocompat-ibility due to its non-corrosive properties, low ion-formationtendency in aqueous environments and a dielectric constantcomparable to that of water [31,62]. The material passivates itselfin vivo by forming of an adhesive oxide layer [31,62]. It alsodisplays the unique property of osseointegration, by which itconnects both structurally and functionally with the underlyingbone [62,75]. Controlled formation of TiO2 and Ti5Si3 on thesurfaces of a number of Ti-alloys can induce apatite formationwhen these surfaces are immersed in a simulated physiologicalmedia of the appropriate ionic concentration, enhancing earlybinding of Ti-alloys to bone [109]. Similarly, treating Ti-alloys withan aqueous solution of NaOH, followed by heat treatment at 500–800 8C, results in a thin titanate layer, which can then form a dense,bone-like apatite layer when placed in physiological media, thusenhancing the strength of the bone/implant interface [75,106]. Forpermanent implants, Ti–6Al–4V has a possible toxic effectresulting from released vanadium and aluminium, so this alloyis being replaced by vanadium- and aluminium-free alloys (Ti–13Nb–13Zr and Ti–12Mo–6Zr) [62].

Table 1 presents some mechanical properties of several metallicbiomaterials. In general, these materials have high tensile andfatigue strength compared with ceramic and polymers. However,the elastic moduli are much higher than that of natural bone,which can cause ‘‘stress shielding,’’ a phenomenon characterisedby bone resorption in the vicinity of implants.

3.1.1. Shape memory alloys

Shape memory alloys (SMA) are materials that retain theiroriginal shape after severe deformations when subjected to heatabove their transformation temperature [24,123,159]. Shapememory alloys have two distinct crystallographic phases, namely,austenite and martensite. The martensitic phase is a low-temperature, stable phase with the absence of stress. The austenitephase is stable at a high temperature and displays a stronger body-centre cubic structure [147]. SMAs are capable of large amounts ofbending and torsional deformation and high strain rates (6–8%) inthe martensitic phase [194]. Another unique property of SMAs ispseudo-elasticity, wherein the two-way phase transformationoccurs at a constant temperature.

The most commonly used SMA is nitinol, an alloy containingapproximately 50% nickel and 50% titanium [89]. Titanium is non-toxic, while nickel is extremely toxic and carcinogenic. However,nitinol forms a passive titanium oxide layer that both acts as a

from corrosion [159]. As illustrated in Fig. 8, nitinol similar to band tendon, has high elasticity, low deformation forces

constant force over wide ranges of strain.Cu-based SMAs, especially Cu–Al–Ni and Cu–Al–Mn, are

commercially available. They have a transformation temperatrange (�200 to 200 8C) similar to nitinol, but higher Youmodulus, better machinability and better stability [89,1though Cu-based alloys have toxic effects [198].

SMAs are used for hard tissue implants in orthopaedics

dentistry due to its porous structure, good mechanical properbiocompatibility and shape memory effect [95,159,181].

3.2. Ceramics

Ceramics are inorganic materials with high compresstrength and biological inertness [59,134,211]. The most comonly used bioceramics are metallic oxides (e.g., Al2O3, Mcalcium phosphate (e.g., hydroxyapatite (HA), tricalcium phphate (TCP), and octacalcium phosphate (OCP)), and glass ceram(e.g. Bioglass, Ceravital) [26,82]. Metallic oxides are considerebe nearly bioinert in biological environments, while calcphosphate and glass ceramics can bond to bone when implanBioceramics have been successfully used for hard tissue replament due to their good biocompatibility and bioactivity. Tbiocompatibility is a direct result of its chemical compositiwhich contain ions commonly found in the physiologenvironment, such as Ca2+, K+, Mg2+, Na+ [211].

Bone tissue becomes integrated into the bioactive ceramthrough the biomineralization of a thin layer of calcium phosphat the interface between the ceramics and the host bony tis[134]. The interstitial body fluid is the very first medium obioactive ceramic interface after being hosted in a bony defect.

structure of the ceramic changes for the biomineralizationcalcium phosphate by the interaction with the body fluid, whcontains various proteins that must be significantly involved wbiomineralization [134,170]. This interfacial layer of calcphosphate is almost independent of the ceramic type. Fig

Fig. 8. Stress vs. strain relationship for superelastic nitinol, stainless steel, bone

tendon [159].

physical barrier to nickel oxidation and protects the bulk material

Table 1Relevant properties of metallic biomaterials [175].

E modulus

[GPa]

Yield strength

[MPa]

Tensile strength

[MPa]

Stainless steel 190 221–1213 586–1351

Co–Cr alloys 210–253 448–1606 655–1896

Titanium 110 485 760

Ti–6Al–4V 116 896–1034 965–1103

Cortical bone 15–30 30–70 70–150Fig. 9. In vitro mechanism of formation of calcium phosphate on the surface of

Na2O–CaO–SiO2 glass in SBS [134].



illuspho

Tappl(HAcharpromimpof thimpis tosuchpolycoattitanas aaltertitanlow

ceraS

phoof ocells[6,1

Sfineweahip

imphighcerawheimplowe[51]surg

BalsoBioadepoenhativeW gbioavertobsesilic[50]vivo

3 pbiom

TableMain

Ca/

1.3

1.5

1.5

1.6

2.0

TableRelev

Alu

Bio

Cal


G Model


Plepr

trates the in vitro mechanism of the formation of calciumsphate on the surface of a Na2O–CaO–SiO2 glass in SBF.he main calcium phosphate materials used for medicalications are indicated in Table 2. Synthetic hydroxyapatite, Ca10(PO4)6(OH)2) is a bioactive material, with chemicalacteristics similar to hard tissues such as bone and teeth, thatotes hard tissue ingrowth and osseointegration when

lanted into the human body [26,211]. The porous structureis material can be tailored to suit the interfacial surfaces of the

lant. As a bulk material, HA lacks sufficient tensile strength ando brittle to be used in most load bearing applications [52]. In

cases, HA is coated onto a metal core or incorporated intomers as composites [28]. HA is frequently used as a bioactiveing on hip prostheses [52]. The ceramic coating on theium implants improves the surface bioactivity but often fails

result of poor ceramic/metal interface bonding [100]. Annative is the production of composite materials containingium and bioceramic as a reinforced phase [100]. Due to thebioresorbability of HA much attention has been paid to TCPmics [28,101].everal in vitro and vivo works have shown that calciumsphates support the adhesion, differentiation and proliferationsteogenesis-related cells (e.g., osteoblasts, mesenchymal stem), besides inducing gene expression in bone cells

79,187,195,210].tructural ceramics like alumina (high purity, polycrystalline,

grained) and zirconia that is toughened and highly resistant tor (TZP and Mg-PSZ) have been used for femoral heads of totalprostheses due to their excellent tribological properties,

roved fracture toughness and reliability [126]. Zirconia has flexural strength and fracture toughness compared to othermics, which makes it more resistant to masticatory forcesn used as crowns with exact precision of fit [104,228]. Zirconialants also accumulate less bacteria in vivo [177] and undergo ar rate of inflammation-associated processes than titanium

. Zirconia has also been used in shoulder reconstructionery and as a coating over titanium in dental implants [53].ioactive glasses, such as Bioglass, and A-W glass–ceramic have

been successfully used for tissue replacement [122,223].ctive glasses stimulate the formation, precipitation andsition of calcium phosphates from physiological solutions,ncing the bone–matrix interface strength. Bioglass is bioac-

with low fracture toughness [211,223], while the bioactive A-lass–ceramic has excellent mechanical properties and highctivity (higher than HA), being clinically used for iliac andebrae prostheses and intervertebral spacers [114]. It wasrved that ionic dissolution products from Bioglass and other

ate-based glasses stimulate gene expression of osteoblasts. Bioactive glasses also stimulate angiogenesis in vitro and in

, as well as antibacterial and inflammatory effects [4,72]. Table

3.3. Polymers

Polymers for medical applications can be naturally derived orsynthetic, the latter of which can be biodegradable or bioinert.Bioinert synthetic materials include polyvinyl chloride (PVC),polyethylene (PE), polypropylene (PP), polymethylmethacrylate(PMMA), polystyrene (PS), polytetrafluoroethylene (PTFE), polye-sters, polyamides (PA-nylon), polyurethanes (PUR), and polysilox-anes (silicone) [26,96]. Biodegradable synthetic polymers includepoly(glycolic acid), poly(lactic acid), their copolymers, and poly(p-dioxanone). Natural polymers include albumin, collagen, cellulose,hyaluronic acid, starch, chitosan, dextran, silk, heparin, and DNA[26,96].

Biodegradable synthetic polymers have been used in a numberof clinical applications, such as resorbable sutures, drug deliverysystems, orthopaedic fixation devices such as pins, rods andscrews, and scaffolds for tissue engineering.

3.3.1. Hydrogels

Hydrogels are cross-linked hydrophilic polymers that exhibitexcellent biocompatibility, causing minimal inflammatoryresponses, thrombosis, and tissue damage. Hydrogels can alsoswell large quantities of water without the dissolution of thepolymer due to their hydrophilic and cross-linked structure, whichgives them physical characteristics similar to soft tissues[97,190,193]. They also present high permeability for oxygen,nutrients, and other water-soluble metabolites. Hydrogels are softand elastic materials, generally used above their glass transitiontemperature (Tg).

Synthetic hydrogels can be formed from poly(ethyleneglycol)(PEG), poly(vinyl alcohol) (PVA), propylene fumarate (PF), poly(-butylene oxide) (PBO), polycrapolactone (PCL), poly(hydroxybut-tyrate) (PHB), polyacylamide, poly(vinyl acetate) (PVamine),poly(vinyl acetate) (PVac,), polyacrylonitrile (PAN), poly(ethyleneoxide) (PEO), poly(propyleneoxide) (PPO), poly(hydroxypropylmethacrylamide) (PHPMA) and poly(2-hydroxyethyl methacry-late) (PHEMA). Biological hydrogels can be formed from hyaluronicacid (HA), alginic acid, agarose, chitin, fibrin, collage, dextran,agarose, gelatine, pullulan and carrageenan.

Hydrogels can be chemically tailored to respond to certainenvironmental stimuli, as so-called temperature-responsive,potential-specific and pH-sensitive gels (Fig. 10) [190]. They areextensively used in medicine for applications such as contact

2 calcium phosphate compounds.

p molar ratio Compound Formula

3 Octacalcium phosphate (OCP) Ca8(HPO4)2(PO4)45H2O

a-Tricalcium phosphate (a-TCP) a-Ca3(PO4)2

b-Tricalcium phosphate (b-TCP) b-Ca3(PO4)2

7 Hydroxyapatite (HA) Ca10(PO4)6(OH)2

Tetracalcium phosphate Ca4(PO4)2O

resents some mechanical properties of several ceramicaterials.

Fig. 10. Smart hydrogels: (A) pH-sensitive hydrogels, (B) temperature sensitive

hydrogels, (C) biomolecule sensitive hydrogels [190].

3ant properties of ceramic biomaterials.

Young’s

modulus [GPa]

Compressive

strength [MPa]

Tensile

strength [MPa]

mina 380 4500 350

glass-ceramics 22 500 56–83

cium phosphate 18–28 517 280–560



lity.icalder

essringandom

g antro-asic

intte,hisin-

igh-s aelt

ows48].hilet to

theylor. Annoticaltiessiteres

ters areith

can


G Model


lenses, biosensors, linings for artificial implants, wound healing,and drug delivery devices [98]. In wound healing applications,hydrogels protect the wound from desiccating, thus providingcontrol of the wound surface hydration, sometimes absorbingexcess exudate and often providing moisture [98].

3.3.2. Shape memory polymers

Shape memory polymers (SMPs) are degradable or non-degradable polymers that allow repeated shape changes andshape retention [153,192,216,221]. At temperatures above Tg thematerial enters a rubbery elastic state, so it can be easily deformedinto any shape. When the material is cooled below its Tg, thedeformation is fixed and the shape remains stable. The originalshape can be recovered by heating the material once again to atemperature higher than Tg [192,216]. SMP presents severaladvantages over SMA, such as the following [192]:

� Lightweight;� The wide range of Tg (from �70 8C to 100 8C) allows a wide

variety of potential applications in different thermal environ-ments;� High shape recovery (up to 400%);� Large reversible changes of elastic modulus between the glassy

and rubbery states;� Excellent biocompatibility;� Easy processing;� Low cost (10% of the cost of existing SMAs).

SMPs are used for orthopaedic and dental applications,bandages and artificial skins, self-tightening suture materials,drug delivery systems, stents, intelligent electrodes and throm-bectomy devices [192,216].

3.4. Composites

A wide range of polymer-based composite materials weredeveloped and investigated for biomedical applications. Thesematerials can be classified as shown in Fig. 11 [172]. A compositematerial made of an avital (non-living) matrix and reinforcementphases is called an ‘‘avital/avital’’ composite. Alternatively, acomposite material comprising a vital (living) and avital (non-living) material is called a ‘‘vital/avital’’ composite [172].

4. Manufacturing processes

Additive electro-chemical and physical processes, throughwhich physical objects are created from computer-generatedmodels, emerged in the 1980s. The basic concept of additivefabrication is layer laminate manufacturing, in which 3Dstructures are formed by laminating thin layers according to 2Dslice data obtained from a 3D model. The main advantages ofadditive electro-chemical and physical techniques are the capacityto rapidly produce very complex 3D models and the ability to usevarious raw materials. When combined with clinical imaging data,these fabrication techniques can be used to produce constructscustomised to the shape of the defect or injury. Some processesoperate at room temperature, thus allowing cell encapsulation and

biomolecule incorporation without significantly affecting viabiThis section describes the most relevant additive electro-chemand physical processes either commercially available or undevelopment.

4.1. Electrospinning

Electrospinning is the most relevant electro-chemical procto produce nano-scale meshes for tissue enginee[1,13,38,48,60,91,133,157,158,167,171,201]. It is a simple

versatile process by which nanofibers with diameters ranging fra few nanometers to several micrometres can be produced usinelectrostatically driven jet of polymer solution (solution elecspinning) or polymer melt (melt electrospinning). The brequirements of an electrospinning apparatus are shownFig. 12 including: (1) a capillary tube with a needle or pipe(2) a high-power voltage supply, and (3) a collector or target. Tcollector can move in the vertical direction, enabling electrospning as an additive technology. Electrical wires connect the hpower supply to both the capillary tube, which containpolymeric solution, and the target. An example of a melectrospinning apparatus is illustrated in Fig. 13. Fig. 14 shtwo different heating configurations for melt electrospinning [Melt electrospinning requires, the polymeric jet to be cooled, wsolution electrospinning relies on the evaporation of the solvenproduce fibres.

Initially, as a result of surface tension, pendant droplets ofsolution are held in place. A conical protrusion, known as a Tacone, is formed when a critical voltage is applied to the systemapproximately straight jet emerges from the cone but it canstand for long. The jet then emerges into a diaphanous and conshape. The conically moving jet experiences bending instabiliand is directed towards the collector, which has the oppoelectrical charge. The solvent evaporates, and dry polymer fibare deposited until the jet reaches the collector. The parameand processing variables affecting the electrospinning processindicated in Table 4. Fig. 15 illustrates meshes produced woptimised and non-optimised parameters. Hollow nanofibers

be prepared by co-axial electrospinning (Fig. 16).

Fig. 12. Electrospinning setup [201].

Fig. 11. Classification of polymer-based composite biomaterials [172]. Fig. 13. Melt electrospinning setup [1].



M

� So

PLtriurCooolikepCecr(g� M

12(P

4.2.

Sobjeinitiusuamaspolya fotwoand

Tgrapfocupho

Fig.

melti

Fig


G Model


Plepr

aterials commonly used in electrospinning include:

lution electrospinning: Polyethylene-co-vinyl acetate (PEVA/A), Polylactic acid (PLA), Polyvinil alcohol (PVA), Polyacryloni-le (PAN), Polycarbonate (PC), Polybenzimidazole (PBI), Poly-ethanes (PU), Nylon6,6 (PA-6,6), Polyethylene oxide (PEO),llagen/PEO, Polymethacrylate (PMMA)/tetrahydroperfluor-ctylacrylate (TAN), Polyaniline (PANI)/Polystyrene (PS), Silk-e polymer with fibronectin functionality, Polyethylene Ter-htalate (PET), Poly vinyl phenol (PVP), Polyvinylchloride (PVC),llulose acetate (CA), PVA/Silica, Poly(2-hydroxyethyl metha-ylate) (HEMA), Polycaprolactone (PCL), PAN/TiO2, PCL/metalold or ZnO), alginate and gelatine.elt electrospinning: HDPE, Polypropylene (PP), Nylon 12 (PA-), Polyethylene terephthalate (PET), Polyethylene naphthalateEN), PET/PEN.

Stereolithography

tereolithographic processes produce three-dimensional solidcts in a multi-layer procedure through the selective photo-ated cure reaction of a polymer [14,15,19–21]. These processeslly employ two distinct methods of irradiation. The first is the

k-based method, in which an image is transferred to a liquidmer by irradiation through a patterned mask. The second usescused UV beam to selectively solidify the liquid resin. These approaches can also be classified into two types, free-surfaceconstrained-surface (Fig. 17) [25,90].wo-photon polymerisation represents a useful stereolitho-hic strategy to produce micro/nanoscale structures bysing femtosecond laser pulses into the volume of a liquidtosensitive polymer or polymer mixture. In this case, the

reactive molecules, which are present in the polymeric mixture,absorb two photons instead of one. The probability of electronicexcitation of a molecule by the simultaneous absorption of twophotons depends quadratically on the incident light intensity[219]. This allows a submicron 3D resolution, in addition toenabling 3D fabrication at a greater depth and an ultrafastfabrication. Lim [132,140] developed the so-called nano-stereo-lithography (NSL), based on the two-photon polymerisationprocess (Fig. 18). The system uses a mode-locked Ti:sapphirelaser beam, with a wavelength of 780 nm, pulse repetition of80 MHz and pulse width less than 100 fs. The beam is scannedacross the focal plane using a set of two galvano mirrors with aresolution of 2.5 nm. Microparts are fabricated using a voxelmatrix scanning method or a contour offset method [132,140].

The main advantages of stereolithographic processes includethe ability to cure quickly at physiological temperatures, enablingthe production of scaffolds for tissue engineering applications.Stereolithographic processes have been used to produce hearingaids, micro needles for transdermal drug delivery and scaffolds fortissue engineering with or without encapsulated cells (Fig. 19)[25,165]. Stereolithography is also used to produce surgical guidesfor the placement of dental implants, temporary crowns andbridges and resin models for lost wax casting [163].

Levy et al. [137] used a direct irradiation stereolithographicprocess to produce hydroxyapatite (HA) ceramic scaffolds fororbital floor prosthesis. A suspension of fine HA powder into a UVphoto-curable resin was formulated and used as building material.The photo-cured resin acts as a binder to hold the HA particlestogether. The resin is then burnt out and the HA powder assemblysintered for consolidation. Similarly, Griffith and Halloran [73]produced ceramic scaffolds using suspensions of alumina, siliconnitride and silica particles with a photo-curable resin. The binderwas removed by pyrolysis and the ceramic structures sintered.Briant and Anseth [32] used a photo-polymerisation process toencapsulate chondrocytes in poly(ethylene oxide) (PEO) hydrogelsstructures with thicknesses varying from 2 to 8 mm. The hydrogelstructures were photo-cured at a low light intensity (�10 mW/cm2) for 10 minutes. The chondrocytes encapsulated in thehydrogel structures and cultured in vitro for 6 weeks remainedviable and produced cartilaginous tissue. The results suggested

14. Two different heating configurations for melt electrospinning: (A) low

ng point polymers, (B) high melting point polymers [48].

Table 4Relevant processing variables affecting the electrospinning process [188].

Parameter Effect on fibre morphology

Applied voltage " Fibre diameter # initially, then "Flow rate " Fibre " (beaded morphologies

occur if the flow rate is too high)

Distance between

capillary and collector "Fibre diameter # (bead morphologies

occur if the distance between the

capillary and collector is too short)

Viscosity " Fibre diameter "Solution conductivity " Fibre diameter # (broad diameter

distribution)

Solvent volatility " Fibres exhibit microtexture (pore on

their surfaces, which increase surface

area)

. 15. Solution electrospinning of PCL. (A) Bead formation due to non-optimised process parameters; (B) mesh produced using optimised processing parameters.



eensizeeenach

toe ofmerlledfold the0 8C2 h,. An

ithpli-GA)ced

cro-ratewasfoldwasingeseith

ad-esinol)/ight


G Model


and the acrylic binder were removed by pyrolysis and the HA grscaffold submitted to a sintering process. The finest channel

achieved was 366 mm with a range of implant porosity betw26% and 52%. Similarly, Kim et al. [107] used an indirect approto produce HA scaffolds. Micro-stereolithography was usedproduce a mould (4.2 mm � 4.2 mm) with an internal pore siz250 mm and a line width of about 350 mm using a liquid polySL5180 resin (Huntsman) (Fig. 20a). Then the mould was fiwith HA nanopowder with a particle size of 500 nm. The scaf(Fig. 20b) was produced through a sintering process. Initially,temperature was increased from room temperature to 110with a heating rate of 5 8C/min, held at this temperature for

and then decreased to room temperature with the same rateanalogous strategy was used by Lee et al. [130].

Lee et al. [131] investigated the effect of 3D scaffolds wembedded growth factor-delivering microspheres for bone apcations. BMP-2-loaded poly(DL-lactic-co-glycolic acid) (PLmicrospheres were incorporated into a 3D scaffold produthrough micro-stereolithography, with a suspension of mispheres and a poly(propylene fumarate) (PPF)/diethyl fuma(DEF) photopolymer. By measuring released profiles in vitro, it

verified that the fabricated microsphere-containing 3D scafcould gradually release the growth factor. The effect of BMP-2

also assessed in vitro by observing cell differentiation usMC3T3-E1 pre-osteoblasts. It was also observed that thscaffolds have a superior bone-regeneration effect compared wscaffolds produced using a conventional method (Fig. 21).

Seck et al. [183] produced porous and non-porous biodegrable hydrogel structures using an aqueous photo-curable rbased on methacrylate-functionalised poly(ethylene glycpoly(D,L-lactide) macromers and Lucirin TPO-L as a visible lFig. 17. Schematic diagram of (A) constrained surface system, (B) free-surface

system [25].

Fig. 16. Co-axial electrospinning setup [158].

Fig. 19. Mushroom-shaped cap of a hearing aid [25].

Fig. 20. (A) SEM image of a mould produced using micro-stereolithography, (B)

internal shape of the HA scaffold [107].

Fig. 18. The nano-stereolithography system [132].

that increasing hydrogel thickness from 2 to 8 mm does not changeeither cell viability or uniformity. Chu et al. [40] developed a lost-mould technique to produce implants with designed channels andconnection patterns. Stereolithography was used to create epoxymoulds designed from negative images of implants. A highlyloaded HA-acrylate suspension was cast into the mould. The mould



phostrucomstrumesstru

AmulplishallowThe

rinsremfabrlaberegion tdem

TreseDevdevesterultrationcontconcreacradiinitipolythe

amoprocfragthe

reaccureconv

Fig. 2Nega

BMP-

stere

Fig.

funct


G Model


Plepr

to-initiator (Fig. 22). After photo-polymerisation, the obtainedctures were extracted with distilled water to remove solublepounds, and dried at ambient conditions for 3 days. Thectures showed good cell seeding characteristics, and humanenchymal stem cells adhered and proliferated well on thesectures.rcaute et al. [7–9] explored stereolithography for fabricating

ti-material, spatially controlled bioactive scaffolds. To accom- multi-material builds, a mini-vat setup was designeding for self-aligning X–Y registration during fabrication.

mini-vat setup allowed the part to be easily removed anded and for different photocrosslinkable solutions to be easilyoved and added to the vat. Multi-material scaffolds wereicated by including controlled concentrations of fluorescentlylled dextran or fluorescently labelled bioactive PEG in differentons of the scaffold. Human dermal fibroblast cells were seededop of the fabricated scaffolds. Spatial control was successfullyonstrated in features as small as 500 mm.o produce multimaterial functionally graded scaffolds,archers from the Centre for Rapid and Sustainable Productelopment of the Polytechnic Institute of Leiria (Portugal) areloping a new stereolithographic fabrication process called the

eo-thermal-lithographic process (STLG). This process usesviolet radiation and thermal energy (produced by IR radia-

) to initiate the polymerisation reaction in a mediumaining both photo- and thermal-initiators (Fig. 23). Theentrations of both initiators are carefully selected, and thetion begins only when there is a particular combination of UV

� Generation of radicals is more efficient;� Small concentrations of the two types of initiators are used,

enabling the radiation to penetrate deeper into the polymer;� Combination of UV radiation and temperature increases the

reaction rate and hence the fractional conversion values;� Curing reaction is more localised, improving the accuracy of the

produced models;� The system has more tunability.

Four subsystems can be considered. Subsystem A uses ultravioletradiation to solidify a liquid resin that contains a certain amount ofphoto-initiator. Subsystem B uses thermal energy produced byinfrared radiation to solidify a liquid resin that contains a certainamount of thermal-initiator. Subsystem C uses heat produced usinginfrared radiation and ultraviolet radiation to solidify a liquid resinthat contains a certain amount of photo-initiator. Subsystem D usesheat produced through infrared radiation and ultraviolet radiationto solidify a liquid resin containing a certain number of thermal-initiators and photo-initiators.

In addition to these key advantages, the system also contains arotating multi-vat that enables the fabrication of multi-materialstructures (Fig. 24). STLG is being developed to produce multi-material microscopic engineering prototypes through nanostruc-tures for exploitation in waveguiding and photonic crystals, multi-material functional graded scaffolds for tissue engineering, otherbiomedical components and micro-functional metallic or ceramicparts.

4.3. Laser sintering and melting processes

Selective laser sintering (SLS) and selective laser melting (SLM)are additive manufacturing processes that use high-energy light

1. Micro-CT images of rat cranial bone at 11 weeks after implantation. (A)

tive control, (B) BMP-2-unloaded particulate leaching/gas foaming scaffold, (C)

2-unloaded micro-stereolithography scaffold, (D) BMP-2-loaded micro-

olithography scaffold [131].

22. Hydrogel structures built by stereolithography using a methacrylate-

ionalised poly(ethyleneglycol)/poly(D,L-lactide) resin [183].

Fig. 23. The micro stereo-thermal-lithographic process: irradiation mechanism.

Fig. 24. The micro stereo-thermal-lithographic process: multi-vat system.

ation and thermal energy [21]. This way, the amount of eachator must be low enough to inhibit the start of themerisation by only one of these two effects. However, atpoint where the two effects intersect each other, sufficientunts of radicals are generated to initiate the polymerisationess. Temperature is used to both produce radicals through thementation of thermal-initiators and simultaneously increaseinitiation and reaction rate of the photo-initiated curingtion. As a result, the extent of cure is increased, and no post-

will be needed. The main advantages of STLG overentional stereolithography are as follows:



up-andthe

ical,eres.urentsherand

byfor

g of23].ousing.

aling


G Model


sources to consolidate powder material [10,47,63,71,110,120,145,225]. Contrary to SLS, SLM uses high-powered laserbeams to directly create 3D metal parts by fusing very fine metallicpowders. Table 5 summarises the major advantages and dis-advantages of SLM to process metal powders.

SLS and SLM have been used to produce both permanent andtemporary implants. Pattanyak et al. [166] explored the use of SLMto produce porous titanium constructs with complicated internalstructures for bone ingrowth applications. The constructs wereproduced using Ti powder of less than 45 mm particle size. Thecompressive strength was in the range of 35–120 MPa when theporosity was in the range of 75–55%. Porous Ti constructs weresubjected to NaOH, HCl, and heat treatment to provide bioactivity.Treated constructs formed bone-like apatite on their surfaces in astimulated body fluid within 3 days. In vivo research also showedthat new bone penetrated into the pores. Similar work was carriedout by Imwinkelried [94] and Wang et al. [212].

Hollander et al. [84] used SLM to produce a wide range of Ti–6Al–4V medical implants, ranging from cylinders with regularporosity to a human vertebra model (Fig. 25). In vitro studies wereperformed with porous structures using human osteoblasts. Cellspreading and proliferation was observed. Similar studies wereperformed at the University of Leuven: some ten different scaffoldgeometries were produced and seeded with human periosteum-derived cells (Fig. 26). After 14 days of culture in a growth medium(GM based on DMEM and bovine serum) and osteogenic medium(OM = GM + dexamethasone + ascorbic acid), the cells were foundto be viable and proliferating in all scaffolds. GM culture resulted inmore cells and a greater extend of pore occlusion.

Partners in Belgium and the Netherlands (universities of Leuvenand Hasselt, AM bureau LayerWise, etc.) cooperated to design andmanufacture complex jaw implants (Fig. 27). The full lowerimplant shown in Fig. 27 was coated with bioceramics andimplanted in the chess of an 83 year old lady. The cavities in theimplant made it to weight only slightly more than a natural jawbone and guaranteed good attachment of muscles and space fornerves.

Hao et al. [76] used SLM to directly process HA and 316Lstainless steel (SS) powder mixture to develop load bearing andbioactive implants. The SS/HA composite implants fabricated usingoptimum parameters exhibited a tensile strength of 95 MPa, whichis adequate for load-bearing applications.

Kruth [119,206,207] explored SLM to produce implant-sported frameworks for dental prostheses using titanium

cobalt-chromium. Quality control has been performed on

produced frameworks to verify their mechanical, chembiocompatible and geometrical properties. The frameworks wclinically tested and are now commonly implanted in patient

SLS and SLS combined with self-propagating high temperatsynthesis (SHS) have also been used to produce Nitinol impla[61,120]. The combined process resulted in implants with highomogeneity in chemical composition, better biocompatibility

more porosity. The surface of porous Nitinol implants madeSHS/SLS had a significantly more favourable structure

mechanically interlocking with bone [185,186]. Laser meltinTi–Ni shape memory alloy was also investigated by Bourell [1

Rimell and Marquis [176] used SLS to produce linear continuUHMWPE solid bodies for clinical applications without pre-heat

Table 5Main advantages and disadvantages of SLM [120].

Advantages Disadvantages

Material No distinct binder and

melt phases

Not suitable for well-

controlled composite

materials (e.g. WC-Co)

Cost and

processing

time

Elimination of time-

consuming and costly

furnace post-treatments

for debinding, infiltration

or post-sintering

High laser power and

good beam quality

(expensive lasers);

smaller scanning

velocities (longer

build times)

Part quality Suitable for producing

fully dense parts in

a direct way

Melt pool instabilities

and higher residual

stresses

Fig. 26. Green fluorescence images revealing live cells (A) and SEM images reve

cell proliferation after 14 days.

Fig. 27. (a) Ti mandible plate and inner scaffold structure produced as one part by

SLM and fitted into model of mandible produced by SLA (stereolithography); (b–d)

full lower jaw in Ti.

Fig. 25. Examples of Ti–6Al–4V parts. (A) Cylinders with cubic pore pattern, (B)

original (right) and analogue (left) human vertebra [84].



ProbwerNylowithingrminusedMedMyKpreopatidelivto pthe

Tskeland

morresuaddirangand

propL

polybindphoinfilbase

Zmicrnanomicrprepa sizCHAprobfabr

HhighspeepowspeescanfragSLS

Awithand

drawinabproctranlasea pofabrtrea

LTechacid117of ththe

proc13–�50800

scafcomtrab


G Model


Plepr

lems related to material shrinkage and material degradatione observed. Das et al. [49] used SLS to produce biomimeticn-6 constructs of cylindrical and cubical periodic geometry 800 mm channels and 1200 mm pillars. Results showed bone

owth into the pore channels after implantation into a Yucatanipig mandible for 6 weeks. SLS of polyamide (PA12) is widely

for individual medical jig manufacturing, e.g., MyKnee1 byacta Switzerland and SurgiGuide1 by Materialise, Belgium. Thenee1 cutting blocks are made to accurately match the surgeonsperative planning and operation, based on an individual

ent’s anatomy. The customised SurgiGuide1 drilling jigs areered together with a LongStop Drill system of bushes and drills

recisely control the drilling depth as planned preoperatively bySimPlant software.he potential of SLS to produce PCL scaffolds for replacement ofetal tissues was demonstrated by research teams of Das [215]

Kruth [205]. Das seeded the PCL scaffolds with bonephogenetic protein-7 (BMP-7) transduced fibroblasts. In vivo

lts show that these scaffolds enhance tissue ingrowth, intion to possessing mechanical properties within the lowere of trabecular bone. The compressive modulus (52–67 MPa)yield strength (2.0–3.2 MPa) were in the lower range oferties reported for human trabecular bone.ee and Barlow [127] coated calcium phosphate powder withmer by spray drying a slurry of particulate and emulsioner. The coated powder was then sintered to fabricate calcium

sphate bone implants. Afterwards, these structures weretrated with calcium phosphate solution or phosphoric acid-d inorganic cement.hou et al. [230] studied the use of bio-nano-compositeospheres consisting of carbonated hydroxyapatite (CHAp)spheres within a PLLA matrix to produce scaffolds. PLLAospheres and PLLA/CHAp nanocomposites microspheres wereared by emulsion techniques. The resulting microspheres hade of 5–30 mm, suitable for the SLS process. The use of PLLA/p nanocomposite microspheres seems to offer a solution to thelem of removing the excess powder from the pores after

ication.ao et al. [78] investigated the use of SLS to fabricate HA mixed-density polyethylene (HDPE) scaffolds. Different scanningds and laser power values were considered. HA and HDPEders with 40% HA by volume ratio were mixed using a high-d blender. The results revealed that for low power or highning speeds the layers were generally not sintered or very

ile. Powder blends of PEEK/HA have also been processed using[199].lthough the SLS processes can build highly complex structures

internal architectures appropriate for bone tissue ingrowthwith the required external geometry, one of the criticalbacks of the existing commercial SLS machines has been their

ility to reach required part-bed temperatures during the directessing of ceramics, which typically have higher glass-sition temperatures [144]. This leads to indirect selectiver sintering where the glass or ceramic particles are mixed withlymeric binder and used as feedstock for the SLS machine toicate the green part. The fabricated green part is later heatted to remove the binder and sinter the ceramic particles.eu’s research group at Missouri University of Science andnology used a silicate based 13–93 bioactive glass with stearic

combining designs such as an inner porous and an outer solid shell,the mechanical strength of the scaffolds could be increased forhuman bone repair applications. Tests with simulated body fluid(SBF) showed a thick layer of HA formation on the surface of thescaffolds after 3 weeks. The in vitro results showed that SLS scaffoldsoffer a rough surface, which is favourable for MLO-A5 cells to attach,grow, and proliferate. The recent in vivo results showed completebridging after 6 weeks in a rat segmental defect model, which wasimplanted with a 13–93 glass SLS scaffold with BMP-2 growth factor[207]. Partial bridging was observed in the control group that did notreceive the BMP-2 growth factor. The results have demonstrated thepotential of SLS in manufacturing scaffolds for low to medium load-bearing applications in human bone repair.

Dalgarno’s research group at Newcastle University used apatitebased bioactive glass–ceramic systems to make scaffolds usingindirect SLS [69,70,217,220]. The desired porosity was achievedafter binder burnout and heat treatment. The scaffolds were laterinfiltrated with phosphate glass to improve their mechanicalproperties. Apatite–mullite scaffolds had flexural strengths of16.2 MPa and those made with apatite–wollastonite had strengthsranging from 35 MPa for a porous part to 102 MPa for a fullyinfiltrated part. Although it was reported that no apatite layerformed when the scaffolds were immersed in SBF, new bone tissuegrowth was observed in the porous structure after 4 weeks ofimplantation in a rabbit tibia.

LENS (Laser Engineered Net Shaping), developed by SandiaNational Laboratories, is an alternative process to SLM. This lasercladding based process uses a CAD-driven, high-power Nd:YAGlaser (recent machines use fibre lasers) focused onto a metalsubstrate to create a molten pool (Fig. 29) [12,77,118]. Metal isthen injected into the melt pool to increase the material volume.An inert gas is often used to shield the melt pool from atmosphericoxygen. Medical device materials commonly processed using LENSinclude titanium alloys, stainless steels, cobalt alloys, shapememory alloys and calcium phosphate bioceramics [12,43,77,118].

Fig. 28. Bioactive glass scaffolds made by SLS process: (A) Green scaffolds with 3D

interconnected straight and helical pores compared to a dense cylinder, (B) sintered

scaffolds with different pore sizes along length and within cross-section, (C)

combination of inner porous, outer hollow structures, (D) SEM image of a fractured

surface showing the degree of densification achieved in the struts of scaffolds made

by SLS.

Fig. 29. Schematic representation of the LENS process [77].

as the binder to make scaffolds using indirect SLS [115–,208]. The experimental results showed that the densificatione struts in the porous scaffold can be improved by controlling

SLS process parameters, particle size, binder content and post-essing schedule. Fig. 28 shows some of the parts made with

93 bioactive glass using SLS. The scaffolds have porosities of% and highly interconnected pores in the range of �300–mm. The compressive strengths varied from �41 MPa for afold that is �50% porous to �157 MPa for a fully dense part. Thepressive strengths of the scaffolds are much higher than aecular bone but not comparable to a human cortical bone. By



thecial

sedthisnts,ionens

ane of.een

as74]

asrge

oralt to

em-nge

ter


G Model


4.4. Electron beam melting process

EBM (Electron Beam Melting) is an additive manufacturingprocess that uses an electron beam to scan a layer of metal powderon a substrate, forming a melt pool [77,113,149,163]. The system(Fig. 30) consists of the electron beam gun compartment and thespecimen-fabrication compartment both kept in a high vacuum.Advantages of using EBM over SLM include very small spot sizes,very high beam-material coupling efficiency, high scanning speedand beam deflection without the use of moving mirrors [77]. Theaccuracy of EBM is in the range of 0.3–0.4 mm, and the surfacefinish tends to be rough, with an Ra value in the range of 25 mm[163]. Table 6 compares EBM and SLM.

EBM has been used to produce titanium root-form implants (Ti–6Al–4V ELI) [37], femur hip implants [79], dental implants [113] andknee replacement implants (Figs. 31 and 32) [162]. Kouja et al. [105]evaluated the in vivo performance of Ti–6Al–4V ELI dental implantsfabricated via EBM and compared it to a commercially availableporous-coated press-fit dental implant (Endopore, Innova Corp,Toronto, Canada). Cylindrical shaped implants 3 mm � 5 mm longwere implanted in a rabbit tibia and retrieved after 6 weekspostoperatively. Histology results revealed osteointegration of

surrounding bone with both implant types, suggesting that

implants produced by EBM perform equally well as commerimplants.

4.5. Extrusion-based processes

The extrusion-based technique, commercially known as FuDeposition Modelling (FDM), was developed by Crump [46]. In

process, objects are formed by thin thermoplastic filamemelted by heating and deposited by a NC controlled extrushead. The material leaves the extruder in a liquid form and hardimmediately. The working platform is contained withinenclosed chamber that is held below the melting temperaturthe thermoplastic material to aid in the bonding process [74]

The use of FDM to produce permanent implants has blimited to the fabrication of anatomic models that can servetemplates for the fabrication of custom implants. Gronet et al. [described the use of FDM to produce 3D models that servetemplates for the fabrication of custom acrylic implants for ladefects with complex contours involving the anterior tempregion or defects where margins are inaccessible or difficuldetect by palpation (Fig. 33).

Extrusion-based processes are widely used to produce tporary implants (scaffolds) for tissue engineering in a wide ra

Fig. 30. Schematic representation of the EBM process [77].

Fig. 31. (A) Co–29Cr–6Mo femoral prototype with mesh structure produced by

EBM, HIP-annealed using ASTM-F75 standard, machine finished and partially

polished, (B) magnified view of the mesh structure section [162].

Table 6Comparison between EBM and SLM [66].

EBM SLM

Thermal source Electron beam Laser

Atmosphere Vacuum Inert gas

Scanning Deflection coils Galvanometers

Energy absorption Conductivity-limited Absorptivity-limited

Powder pre-heating Use electron beam Use infrared heaters

Scan speed Very fast, magnetically

driven

Limited by galvanome

inertia

Energy costs Moderate High

Surface finish Moderate to poor Excellent to moderate

Feature resolution Moderate Excellent

Materials Metals (conductors) Metals, ceramics and

polymers

Fig. 33. (A) FDM anatomic model revealing posterior fossa defect along foramen

magnum, difficult to restore by conventional methods, (B) acrylic implant on FDM

anatomic model restoring oval aperture of foramen magnum, (C) scalp closure after

model implantation [74].Fig. 32. Ti–6Al–4V ELI acetabular cups.



of pexplextrrapirigiddiamadju

Tengimacscafload25 8comcom5.11doubsolustiffnscaffconnof sc

RPEOand

to thW

DepconttionprocwhematProlfibroand

plasW

Fibrpolyfoldartiction

Fig. 3with


G Model


Plepr

olymers and polymer/ceramic composites. Koh et al. [111]oited the fact that when a warm PCL-HA/acetone solution isuded into a reservoir containing ethanol, the extruded filamentdly solidifies via solvent extraction, producing a continuous

filament to fabricate macro-channelled scaffolds. Theeter and morphology of the filament were controlled by

sting the deposition speed and volume flow rate.ellis et al. [200] used micro CT to create biomimetic tissueneering scaffolds. CAD models were exported to an FDMhine, producing polybutylene terephthalate (PBT) trabecularfolds. The scaffolds were compression tested at two different

rates (49 N/s and 294 N/s). Some scaffolds were soaked in aC saline solution for 7 days before compression. Whenpressed at 49 N/s, the dry trabecular scaffolds had apressive stiffness ranging from 2.46 � 0.55 MPa to

� 1.89 MPa. At 294 N/s, the compressive stiffness values roughlyled. It was also observed that soaking the scaffolds in saline

tion had an insignificant effect on stiffness and that compressiveess decreased as pore size increased. Compressive trabecular

olds matched bone samples in porosity. However, physiologicectivity density and trabecular separation requires optimisationaffold processing.agaert et al. [169] used a Bioscaffolder to extrude PCL, PCL-, PCL-Collagen and PLA scaffolds mimicking heart valve leafletsarteries, and compared their elastic and mechanical propertiesose of natural tissues.ang et al. [209] used a process called Precision Extruding

osition (PED) to directly fabricate PCL scaffolds with arolled pore size of 250 mm and designed structural orienta-s (08/908, 08/1208 or combination of both patterns). In thisess, material in pellet or granule form is fed into a chamber,re it is liquefied. Pressure from a rotating screw forces theerial down a chamber and out through a nozzle tip.iferation studies were performed using cardiomyoblasts,blasts and smooth muscle cells. The surface hydrophilicitytotal surface energy of PCL scaffolds was also increased withma treatment.

oodfield et al. [218] used an FDM-like technique, called 3De Deposition, to produce poly(ethylene glycol)-terephthalate-(butylenes terephthalate) (PEGT/PBT) block co-polymer scaf-s with a 100% interconnecting pore network for engineeringular cartilage (Fig. 34). By varying the co-polymer composi-, porosity and pore geometry, scaffolds were produced with a

range of mechanical properties closely resembling articularcartilage. The scaffolds seeded with bovine chondrocytes sup-ported a homogeneous cell distribution and subsequent cartilage-like tissue formation.

Researchers from Tsinghua University (China) developed aprocess called Low-temperature Deposition Manufacturing (LDM)to produce scaffolds in low-temperature environments under 0 8C[222]. The LDM system comprises a multi-nozzle extrusion processand a thermally induced phase separation process. Scaffoldshaving a macroporous structure larger than 100 mm in diameterand a microporous structure smaller than 100 mm have beenreported. The LDM process was used to produce poly(L-lactide)(PLLA) and TCP composite scaffolds with BMP growth factor. Thescaffolds were implanted into rabbit radius and canine radius withlarge-segmental defects. After 12 weeks, it was possible to observethat the rabbit radius defect was successfully repaired, and theregenerated bone had properties similar to the healthy bone.

Moroni et al. [160] reported a strategy to create hollow fibreswith controlled shell thicknesses and lumen diameters, organisingthem into 3D scaffolds. Hollow fibres (Fig. 35), were produced byextruding a blend of poly(butylmethacrylate-methylmethacrylate)(P(BMA/MMA)) and poly(ethylene oxideterephtalate)-co-poly(bu-tylene terephtalate) (PEOT/PBT) using the Bioplotter system. Whileflowing through the nozzle of the extruder, due to viscositydifferences, the polymer with lower viscosity tends to shifttowards the walls. The consequent separation of the polymersproduces a stratification effect. Hollow fibres are produced byremoving the core polymer by selective dissolution. It was alsoobserved that bovine primary articular chondrocytes grow andform ECM not only in the scaffold macropores but also inside thehollow cavities. The use of these hollow matrices for selective drugrelease is being investigated.

The Centre for Rapid and Sustainable Product Development ofthe Polytechnic Institute of Leiria (Portugal) developed anextrusion-based system for scaffold fabrication called BioExtruder.This is a highly reproducible and low cost system enabling thecontrolled definition of pores into the scaffold to modulatemechanical strength and molecular diffusion, as well as thefabrication of multi-material scaffolds [58]. It comprises twodifferent deposition systems: one a rotational system for multi-material deposition achieved using a pneumatic mechanism andthe other for a single material deposition that uses a screw tofacilitate the deposition process (Fig. 36). The rotational system hasfour reservoirs, two with temperature control and two without. Alarge number of nozzle diameters ranging from 0.1 to 1 mm can beused. PCL, PCL/HA, PCL/TCP, PCL/graphene and PCL/PLA were the

4. (A) 3D Fibre Deposition system, (B) SEM sections of 3D deposited scaffolds

varying deposition geometries [218].

Fig. 35. SEM micrographs of a scaffold (A) before and (B) after leaching out the core

material [160].



ogyous

toTheastetionmmThecedth a

theing

tingtantzed

ieve therce

romthe

uldngeare

sive ofing

ed a theriorF inad-

, inuct

thatgesualr or3Dereareoneare

ctorvice

to be

, (D)

most


G Model


materials selected to produce porous scaffolds. Chemical, mor-phological, and in vitro biological evaluation performed on thepolymeric constructs revealed the BioExtruder’s high potential forproducing 3D scaffolds with regular and reproducible macroporearchitectures, without inducing relevant chemical and biocompat-ibility alterations of the material [58]. Several control parameters,such as temperature, screw rotation velocity, deposition velocityand slice thickness, as well as their direct influence on themorphological and mechanical properties of the extruded scaf-folds, were studied [57]. Experimental results reveal that thedeposition velocity and the screw rotation velocity have thehighest influence in terms of the filament diameter and as aconsequence on the porosity and mechanical behaviour of thestructures [57].

Extrusion based processes have also been used to buildscaffolds using ceramics and, more recently, bioactive glasses.Cesarano et al. [34,35] developed a process called Robocasting atSandia National Laboratories (USA). In this process, the paste isextruded through an orifice while the table moves relative to theorifice. The process has been used to make scaffolds based on HAslurries [36]. The parts made were later heat treated to sinter theHA particles. The mechanical properties of the HA based scaffoldswere comparable to human cortical bone with appropriate poresizes and porosities. To overcome limitations in the process inorder to build customised and anatomically shaped implants, anoversized scaffold was fabricated and later machined so as to fit inthe mandible region of a 73-year old female patient as a proof ofconcept. The preparation of colloidal inks and the technique itselfin preparing 3D periodic lattice structures were further researchedby Lewis et al. [138,139,191]. Recently, this technique was used byFu et al. [65] to make bioactive lattice structures for bone tissueengineering. In this process, a bioactive glass based ink is loaded inthe syringe and the lattice is printed on an alumina substrate insidea non-wetting oil reservoir. Compressive strengths of �136 MPawere reported for �60% porous scaffolds.

Researchers at Missouri University of Science and Technol(USA) have developed a process of extruding and depositing aquebased paste called Freeze-form Extrusion Fabrication (FEF)produce lattice structures in a freezing environment [54].

fabrication technique utilises a 3D gantry system. The pextrusion nozzle movement is facilitated in X, Y and Z direcusing orthogonal linear slides, which can cover a distance of 250

and is controlled by a programmable multi-axis controller.

extruder ram is connected to a syringe (paste container) for forpaste extrusion. The syringe is enclosed in a heating sleeve witemperature controller to prevent the paste from freezing insyringe as the entire setup is encased in a freezer box, where freeztemperatures down to �30 8C can be achieved by means of ejecliquid nitrogen. 13–93 bioactive glass powder, binder, surfacand dispersant are mixed in different quantities with de-ioniwater to form a semi-solid paste with a specific viscosity to achuniform extrusion through a nozzle. The signal generated fromload cell is used for feedback control to provide a constant ram foduring the extrusion process. Fig. 37 shows the FEF operation fthe CAD model to the final sintered part, as well as

microstructure of the sintered part.The bioactive glass scaffolds made using the FEF process co

achieve porosities in the range of �50% and pore sizes in the raof �500 mm. The mechanical properties of the scaffolds

comparable to that of a human cortical bone with compresstrengths measuring �140 MPa with an elastic modulus�5.5 GPa. The scaffolds were incubated for 6 days after havbeen seeded with MLO-A5 cells [88]. The MTT test results showgood amount of proliferation of metabolically active cells onscaffolds. These cells covered almost the entire surface and intepores of the scaffold, demonstrating the strong potential of FEfabricating ceramic and glass based biomedical implants for lobearing applications.

Biocell Printing is a multi-head extrusion-based systemdevelopment at the Centre for Rapid and Sustainable ProdDevelopment of the Polytechnic Institute of Leiria (Portugal)

enables the integration and synchronisation of the different staof production and culture of 3D matrices with reduced manintervention [18]. Depending on the chosen strategies (acellulacellular scaffolds), a precision robotic arm transfers the

scaffolds between the construction area (zone 1) to zone 2, whthey are sterilised (Fig. 38). After sterilisation, scaffolds

homogenously seeded with cells using a robotic dispenser (z3). Finally, 3D constructs with embedded or seeded cells

cultured in vitro under dynamic conditions in the biorea(zone 4). The integration of the different stages into a single de

Fig. 36. (A) Bioextruder system, (B) single-material extrusion, (C) multi-material

extrusion system.

Fig. 37. 3D lattice structure creation by FEF process (A) CAD model of the part

built, (B) the FEF machine set-up, (C) fabricated scaffold after sintering

microstructure of the polished cross-section showing good bonding and al

complete densification of the struts.

Fig. 38. Schematic representation of the Biocell Printing system; zone 1 – scaffold fabrication; zone 2 – sterilisation; zone 3 – cell deposition; zone 4 – cell culture in bioreactor.



signprod

Mthreeleclayewithwering

werthicexpeshowscafand

colo

4.6.

Stagetechbon

Tthrestrepowformpowthenunbstreup mordi

KcreapowThe

scafsizewer

Tprolshap(diasalt

diffedistrwithmus

CdemhumsoluThe


G Model


Plepr

ificantly reduces the risk of contamination and increasesuctivity and the possibility of direct clinical application.ota et al. [161] developed dual-scale scaffolds consisting of

e-dimensional constructs of aligned PCL microfilaments andtrospun PLGA fibres (Fig. 39). PCL constructs composed byrs of parallel microsized filaments (08/908 lay-down pattern),

a diameter of 365 mm and interfilament distance of 191 mm,e produced using a melt extrusion-based additive manufactur-technique. PLGA electrospun fibres with a diameter of �1 mme collected on top of the PCL constructs with differentknesses, showing a certain degree of alignment. Cell cultureriments employing the MC3T3 murine preosteoblast cell lineed good cell viability and adhesion on the dual-scale

folds. The influence of electrospun fibres on cell morphologybehaviour was evident, creating a structural bridging for cellnisation.

Inkjet printing processes

everal additive manufacturing processes have taken advan- of ink-jet technology to build 3D parts. Inkjet printingnology can be summarised by two main configurations: ading method and a build-up method.he bonding method was developed at MIT (USA) and is callede-dimensional printing (3D Printing). The process deposits aam of microparticles of a binder material over the surface of ader bed, joining particles together where the object is to beed. A piston lowers the powder bed so that a new layer ofder can be spread over the surface of the previous layer and

selectively joined to it. After the building process, theounded powder is removed, and the porous model must bengthened by a conventional pre-sintering process. The build-

ethod emits a stream of binder microparticles to an exact co-nate.im et al. [108] employed 3DP with particulate leaching tote porous scaffolds, using polylactide-coglycolide (PLGA)der mixed with salt particles and a suitable organic solvent.salt particles were leached using distilled water. Cylindrical

folds measuring 8 mm (diameter) by 7 mm (height) with pores of 45–150 mm and 60% porosity were fabricated. Hepatocytese successfully attached to the scaffolds.he influence of pore size and porosity on cell adhesion andiferation were investigated by Zeltinger et al. [229]. Disced poly(L-lactic acid) (L-PLA) scaffolds measuring 10 mm

within fibrin channels forming a tubular lining. An alternativeprocess has been developed by Mironov et al. [154–156] and Yanet al. [226], who developed the concept of cell printing. Thisprocess prints gels, single cells and cell aggregates offering apossible solution for organ printing [151,154–156,224,226]. To beused for cell printing, the thermal or piezo-tip printheads and inkcartridges are modified to allow bioinks to be printed [95]. Thesebioinks usually consist of aqueous media, thermoreversiblepolymers, or polymer/hydrogel precursors combined with livingcells [178]. Laser-assisted cell-printing techniques have also beendeveloped [151,178]. These techniques comprise the so-calledlaser guidance direct write (LG DW), laser-induced forwardtransfer and modified laser-induced forward transfer processes(Fig. 40) [87,178]. The LG DW process was the first reportedtechnique to print viable cells by forming patterns of embryonic-chick spinal-cord cells on a glass slide. Shortly after this, modifiedlaser-induced forward transfer techniques (LIFT) and modifiedinkjet printers were also used to print viable cells and proteins,followed by the recently introduced electrohydrodynamic jetting(EHDJ) method [42,178].

4.7. Summary

Table 7 summarises the actual state of the art of biomaterialprocessing with appropriate electro-chemical and physical addi-tive manufacturing for the fabrication of both temporary andpermanent implants. However, this is an extensively researchedarea, and new options both in terms of materials and processes arecontinuously emerging.

5. Surface treatments and coating

An important requirement for the clinical success of implants is astrong and effective connection between the implant and the tissue[173]. Surface roughness has been suggested as one important factorfor establishing clinically reliable bone attachments [30,152]. In thecase of scaffolds, surface roughness plays an important role inadhesion, proliferation, differentiation and overall cell viability [39].If the roughness level is too high, cells will not be able to establish

Fig. 39. Fabrication strategy to produce dual-scale polymeric scaffolds [161].

Fig. 40. Schematic representation of the modified laser-induced forward transfer

process [178].

meter) by 2 mm (height) were produced through both 3DP andand leaching methods. The scaffolds were produced with tworent porosities (75% and 90%) and four different pore sizeibutions (<38, 38–63, 63–106 and 106–150 mm), and tested cell cultures using canine dermal fibroblasts, vascular smoothcle cells and microvascular epithelial cells.ui et al. [47] used a modified thermal inkjet printer andonstrated the feasibility of printing microvasculature withan microvascular endothelial cell suspension in thrombintions onto fibrinogen solutions, which served as the substrate.printed cells achieved the capacity to interact and proliferate



y ofan-sed, orine,ups

theThetionins,

ates3].

and

cy

0

g;


G Model


interconnections, which will dramatically compromise prolifera-tion/migration. Experiments performed with compact carbon nano-tubes revealed a higher number of osteoblasts adhering to thesurface of the nano-tubes when compared with fibroblasts,chondrocytes and smooth muscle cells. Despite not yet being wellunderstood, it is thought that nano-scale topology modulates theinterfacial forces between the cells and the implants.

Polymer surface modification is a useful tool to improve thescaffold biofunctionality, creating or increasing specific binding siteswhere bioactive ligands may be immobilised to regulate specificcellular responses [125,168]. Plasma modification processes havebeen increasingly used to modify material surface properties due to

their ability to tune, in a controlled way, the surface densitdifferent functional groups without altering the implant’s mechical properties [23,189]. Independently of the plasma process u(grafting with non-polymerizable gases such as O2, N2, NH3

polymerisation with organic monomers such as allyl amacrylamide, acrylic acid), scaffolds displaying surface polar gro(e.g., NH2, COOH, OH, etc.) may alone increase the bioactivity ofsubstrate and improve cell adhesion and proliferation [189].

plasma processes can also be quite helpful for the immobilisaand deposition of biomolecules such as enzymes, peptides, protepolysaccharides and others, onto plasma-modified substrdisplaying properly selected binding functional groups [143,19

Table 7Actual state-of-the-art of biomaterials processing with appropriate electro-chemical and physical additive manufacturing for the fabrication of both temporary

permanent implants.

Principle Materials state Material class Application Accura

(mm)

Liquid Solid

Electrospinning Polymer melts or

polymeric solutions

Naturally derived and

synthetic polymers

including SMPs

Temporary implants 0.1

Stereolithography Photo-sensitive

polymers

Naturally derived and

synthetic polymers;

polymeric systems highly

reinforced with metallic

and ceramic powders;

hydrogels; SMPs, vital/

avital composites

Temporary implants; drug delivery

systems; functional graded scaffolds;

scaffolds encapsulating cells; scaffolds

containing growth-factors to be used as

controlled-release systems

Permanent constructs can be produced

through a binder removal and sintering

process of highly reinforced liquid

polymeric systems with metallic

particles

0.5–50

Laser sintering Polymer, coated ceramic and

metallic powders or ceramic/

polymer and metal/polymer

powders blends

Synthetic polymers,

metals including SMAs,

ceramics including bioglass

Permanent and temporary implants;

scaffolds containing growth-factors to

be used as controlled-release systems

50

Laser melting Polymeric, metallic and

ceramic powders

High melting point synthetic

polymers, metals including

SMA, ceramics

Permanent implants 20

Electron-beam melting Metallic powder Metals Permanent implants 200<

Extrusion-based Polymer, polymer/ceramic

powders and filaments

Naturally derived polymers,

synthetic polymers,

hydrogels, polymer/

ceramic composites

Extensively used to produce temporary

implants

Fabrication of masters for the indirect

fabrication of permanent implants

It is possible to incorporate

growth-factors and fabricate of

controlled-release scaffolds

100

Inkjet printing

Bonding method Polymer, metal and

ceramic powders

Naturally derived polymers,

synthetic polymers, polymer/

ceramic composites

Fabrication of temporary implants

Fabrication of masters for the indirect

Fabrication of both permanent and

temporary (ceramics) implants

50

Build up method Liquid polymers Hydrogels, vital/avital

composites

Scaffolds incorporating cells, proteins

and growth-factors

20–10

Table 8Techniques to deposit bioresorbable coatings of calcium orthophosphates on metal implants [196,227].

Technique Thickness Advantages Disadvantages

Thermal spraying 30–200 mm High deposition rates; low cost Line of sight technique; high temperatures induce

decomposition; rapid cooling produces

amorphous coatings

Sputter coating 0.5–3 mm Uniform coating thickness on flat substrates;

dense coating

Line of sight technique; expensive; time consumin

produces amorphous coatings

Pulsed laser deposition 0.05–5 mm Coating by crystalline and amorphous phases;

dense and porous coating

Line of sight technique

Dip coating 0.05–0.5 mm Inexpensive; coatings applied quickly; can

coat complex substrates

Requires high sintering temperatures; thermal

expansion mismatch

Sol–gel technique <1 mm Can coat complex shapes; low processing

temperatures; relatively inexpensive as

coatings are very thin

Some processes require controlled atmosphere

processing; expensive raw materials

Electrophoretic deposition 0.1–2.0 mm Uniform coating thickness; rapid deposition

rates; can coat complex substrates

Difficult to produce crack-free coatings; requires

high sintering temperatures

Hot isostatic pressing 0.2–2.0 mm Produces dense coatings Cannot coat complex substrates; high temperature

required; thermal expansion mismatch; elastic

property differences; expensive; removal/interaction

of encapsulation material

Electrochemical deposition 0.05–0.5 mm Uniform coating thickness; rapid deposition

rates; can coat complex substrates;

moderate temperature, low cost

The coating/substrate bonding is not strong enough



Tblastiontioncondprodthe

the

appoimpbiorimp

6. C

RphycustbiocmetSMPcombiolHowprocgrowoppinclu

� Esas� Ap

effi� St� Ac

prhe� De

tis� En

facpr� Pa

de� Sc

to

Ack

ACIRPmanedgeScieEME

Refe

[1]

[2]

[3]

[4]

[5]

[6]


G Model


Plepr

he surfaces of metallic implants can be modified by coatings,ting with various substances, acid treatments, or a combina- of such treatments [196,227]. Generally, chemical composi-, charge, and tension of an implant surface are criticalitions for cell response. Acid etching will, in most cases,uce an increased thickness of the surface oxide layer and alter

physical and chemical properties. It has been suggested thatchemical properties of the oxide layer may benefit bonesition, but its thickness and microstructure are of less

ortance. Table 8 presents relevant techniques to depositesorbable coatings of calcium orthophosphates on metallants.

onclusions and future challenges

ecent investigations with additive electro-chemical andsical processes demonstrated the potential to fabricateomised permanent and temporary implants. A wide range ofompatible materials is available, ranging from metals andallic alloys to ceramics and polymers, including hydrogels ands. Several techniques also show potential for processingposite materials that combine synthetic materials andogical ones, such as cells, proteins and growth factors.ever, the use of additive electro-chemical and physicalesses in the medical field are still in its early life. The rapidlying field of biomanufacturing faces significant challenges and

ortunities. Relevant challenges to be addressed in the futurede:

tablishing a directory materials and related processes andsembly techniques.plying both nano and micro technologies for enhancingcacy and precision.

andardising processes, design and metrology tools.hieving a fundamental understanding of manufacturingocesses and convergence of techniques for best and affordablealth care.velopment of in situ manufacturing strategies, such as in situ

sue engineering.hancing multidisciplinarity, linking clinicians and engineers toilitate further developments and the clinical translation of the

oducts/systems being investigated.ckaging, handling, transportation and accurate tracking, andploying of biomanufactured parts and their building blocks.aling up additive electro-chemical and physical processeswards clinical application.

nowledgements

uthors wish to thank distinguished members of STC-E and, at large, for valuable suggestions in the preparation of thisuscript. Paulo Bartolo and Gideon Levy also like to acknowl-

the support provided by the Portuguese Foundation fornce and Technology through the Strategic Project (PEST-OE//UI4044/2011).

rences

Agrawal S, Wendorff JH, Greine A (2008) Use of Electrospinning Technique for

[7] Arcaute K, Mann B, Wicker R (2010) Stereolithography of Spatially ControlledMulti-material Bioactive Poly(ethylene glycol) Scaffolds. Acta Biomaterialia6:1047–1054.

[8] Arcaute K, Mann BK, Wicker RB (2006) Stereolithography of Three-dimen-sional Bioactive Poly(ethyleneglycol) Constructs With Encapsulated Cells.Annals of Biomedical Engineering 34:1429–1441.

[9] Arcaute K, Mann BK, Wicker RB (2011) Practical Use of Hydrogels in Stereo-lithography for Tissue Engineering Applications. in Bartolo PJ, (Ed.) Stereo-lithography: Materials, Processes, Applications, Springer.

[10] Averyanova M, Bertrand P, Verquin B (2011) Manufacture of C0–Cr DentalCrowns and Bridges by Selective Laser Melting Technology. Virtual andPhysical Prototyping 6:179–185.

[11] Badylak S, Freytes DO, Gilbert TW (2009) Extracellular Matrix as a BiologicalScaffold Material: Structure and Function. Acta Biomaterialia 5:1–13.

[12] Bandyopadhyay A, Krishna BV, Xue W, Bose S (2009) Application of LaserEngineered Net Shaping (LENS) to Manufacture Porous and FunctionallyGraded Structures for Load Bearing Implants. Journal of Materials ScienceMaterials in Medicine 20:S29–S34.

[13] Barnes CP, Sell SA, Boland ED, Simpson DG, Bowlin GL (2007) NanofiberTechnology: Designing the Next Generation of Tissue Engineering Scaffolds.Advanced Drug Delivery Reviews 59:1413–1433.

[14] Bartolo PJ (2007) Photo-curing Modeling: Direct Irradiation. InternationalJournal of Advanced Manufacturing Technology 32:480–491.

[15] Bartolo PJ (2011) Stereolithographic Processes. in Bartolo PJ, (Ed.) Stereo-lithography: Materials, Processes, Applications, Springer.

[16] Bartolo PJ, Almeida HA, Rezende RA, Laoui T, Bidanda B (2008) AdvancedProcesses to Fabricate Scaffolds for Tissue Engineering. in Bidanda B, BartoloPJ, (Eds.) Virtual Prototyping & Bio-manufacturing in Medical Applications,Springer.

[17] Bartolo PJ, Chua CK, Almeida HA, Chou SM, Lim ASC (2009) Biomanufacturingfor Tissue Engineering: Present and Future Trends. Virtual and PhysicalPrototyping 4:203–216.

[18] Bartolo PJ, Domingos M, Gloria A, Ciurana J (2011) Biocell Printing: IntegratedAutomated Assembly System for Tissue Engineering Constructs. CIRP Annals –Manufacturing Technology 60:271–274.

[19] Bartolo PJ, Gaspar J (2008) Metal Filled Resin for Stereolithography MetalPart. CIRP Annals – Manufacturing Technology 57(1):235–238.

[20] Bartolo PJ, Lenz E (2006) Computer Simulation of Stereolithographic CuringReactions: Phenomenological Versus Mechanistic Approaches. CIRP Annals –Manufacturing Technology 55(1):221–225.

[21] Bartolo PJ, Mitchell G (2003) Stereo-thermal-lithography. Rapid PrototypingJournal 9:150–156.

[22] Bartolo PJ, Chua CK (2008) Editorial: Celebrating the 70th Anniversary ofProfessor Yongnian Yan: A life Dedicated to Science and Technology. Virtualand Physical Prototyping 3:189–191.

[23] Bazaka K, Jacob MV, Crawford RJ, Ivanova EP (2011) Plasma-assisted SurfaceModification of Organic Biopolymers to Prevent Bacterial Attachment. ActaBiomaterialia 7:2015–2028.

[24] Bellouard Y (2008) Shape Memory Alloys for Microsystems: A Review from aMaterial Research Perspective. Materials Science and Engineering A StructuralMaterials Properties Microstructure and Processing 481:582–589.

[25] Bertsch A, Renaud P (2011) Microstereolithography. in Bartolo PJ, (Ed.)Stereolithographic Processes, Stereolithography: Materials, Processes, Applica-tions, Springer.

[26] Binyamin G, Shafi BM, Mery CM (2006) Biomaterials: A Primer for Surgeons.Seminars in Pediatric Surgery 15:276–283.

[27] Bittner B, Mader K, Kroll C, Borchert HH, Kissel T (1999) Journal of ControlledRelease 59:23–32.

[28] Bonfield W (1988) In Vivo Evaluation of Hydroxyapatite Reinforced Poly-ethylene Composites. in Ducheyne P, Lemons JE, (Eds.) Materials Character-istics vs. In Vivo Behaviour, New York Academy of Science.

[29] Boretos JW, Eden M (1984) Contemporary Biomaterials, Material and HostResponse, Clinical Applications, New Technology and Legal Aspects, Noyes Pub-lications, Park Ridge, NJ.

[30] Boyan BD, Hummert TW, Dean DD, Schwartz Z (1996) Role of MaterialSurfaces in Regulating Bone and Cartilage Cell Response. Biomaterials17:137–146.

[31] Breme HJ, Biehl V, Helsen JA (1998) Metals and Implants. in Helsen JA, BremeHJ, (Eds.) Metal as Biomaterials, Wiley.

[32] Bryant S, Anseth KS (2001) The Effects of Scaffold Thickness on TissueEngineered Cartilage in Photocrosslinked Poly(ethyleneoxide) Hydrogels.Biomaterials 22:619–626.

[33] Carr BC, Goswami T (2009) Knee Implants – Review of Models and Biome-chanics. Materials and Design 30:398–413.

[34] Cesarano, J., Calvert, P., 2000, Freeforming Objects With Low-binder Slurry,U.S. Pat. 6027326.

[35] Cesarano J, Segalman R, Calvert P (1998) Robocasting Provides MoldlessFabrication from Slurry Deposition. Ceramic Industry 148:94–102.

Biomedical Applications. Polymer 49:5603–5621. Agrawal CM, Huang D, Schmitz JP, Athanasiou KA (1998) Elevated Tempera-

ture Degradation of 50:50 of PLA-PGA. Tissue Engineering 3:345–352. Agrawal CM, McKinney JS, Lanctot D, Athanasiou KA (2000) Effects of Fluid

Flow on the In Vitro Degradation Kinetics of Biodegradable Scaffolds forTissue Engineering. Biomaterials 21:2443–2453.

Allan I, Newman H, Wilson M (2001) Antibacterial Activity of ParticulateBioglass against Supra- and Subgingival Bacteria. Biomaterials 22:1683–1687.

Andrews KD, Hunt JA, Black RA (2007) Effects of Sterilisation Method onSurface Topography and In Vitro Cell Behaviour of Electrostatically SpunScaffolds. Biomaterials 28:1014–1026.

Annaz B, Hing KA, Kayser M, Buckland T, Di Silvio L (2004) Porosity Variationin Hydroxyapatite and Osteoblast Morphology: A Scanning Electron Micro-scopy Study. Journal of Microscopy 215:100–110.


[36] Cesarano J, Dellinger JG, Saavedra MP, Gill DD (2005) Customization of Load-bearing Hydroxyapatite Lattice Scaffolds. International Journal of AppliedCeramic Technology 2:212–220.

[37] Chahine G, Koike M, Okabe T, Smith P, Kovacevic R (2008) The Designand Production of Ti–6Al–4V ELI Customized Dental Implants. JOM 60:50–55.

[38] Chronakis IS (2005) Novel Nanocomposites and Nanoceramics Based onPolymer Nanofibers Using Electrospinning Process – A Review. Journal ofMaterials Processing Technology 167:283–293.

[39] Chu PK, Chen JY, Wang LP, Huang N (2002) Plasma-surface Modification ofBiomaterials. Materials Science and Engineering R 36:143–206.

[40] Chu T-MG, Halloran JW, Hollister SJ, Feinberg SE (2001) HydroxyapatiteImplants with Designed Internal Architecture. Journal of Materials ScienceMaterials in Medicine 12:471–478.



ssesneer-

reo-

anialal of

mi K,ium36.

of alant

1. andlica-

aserate-

IIHArialsteri-

n JA,ate)/

26:

rican

tific

cium877.li H-Juallyrials

on ofints.

31.nces620.011)Free

tors.

ticalional

mersites.

e Artater-icine

sicalrs in

rtoloions,

oam.

nicalation

ices,Poly-

l and47.of aS11.

stedtPhy-


G Model


[41] Clare AT, Chalker PR, Davies S (2008) Selective Laser Melting of HighAspect Ratio 3D Nickel–titanium Structures Two Way Trained for MEMSApplications. International Journal of Mechanics and Materials in Design4:181–187.

[42] Colina M, Serra P, Fernandez-Pradas JM, Sevilla L, Morenza JL (2005) DNADeposition through Laser Induced Forward Transfer. Biosensors and Bioelec-tronics 20:1638–1642.

[43] Comesana R, Lusquinos F, del Val J, Malot T, Riveiro A, Quintero F, BoutinguizaM, Aubry P, Pou J (2010) Three-dimensional Laser-assisted Processing ofBioceramics. Physics Procedia 5:193–201.

[44] Cottam E, Hukins DWL, Lee K, Hewitt C, Jenkins MJ (2009) Effect of Sterilisa-tion by Gamma Irradiation on the Ability of Polycaprolactone (PCL) to Act asScaffold Material. Medical Engineering & Physics 31:221–226.

[45] Crapo PM, Gilbert TW, Badylak SF (2011) An Overview of Tissue and WholeOrgan Decellularization Processes. Biomaterials 32:3233–3243.

[46] Crump, S.S., 1989, Apparatus Method for Creating Three-dimensionalObjects, U.S. Pat. 5121329.

[47] Cui X, Boland T (2009) Human Microvasculature Fabrication using ThermalInkjet Printing Technology. Biomaterials 30:6221–6227.

[48] Dalton PD, Grafahrend D, Klinkhammer K, Klee D, Moller M (2007) Electro-spinning of Polymer Melts: Phenomenological Observations. Polymer 48:6823–6833.

[49] Das S, Hollister SJ, Flanagan C, Adewunmi A, Bark K, Chen C, Ramaswamy K,Rose D, Widjaja E (2002) Computational Design, Freeform Fabrication andTesting of Nylon-6 Tissue Engineering Scaffolds. Rapid Prototyping Journal3(5):205–210.

[50] Day RM (2005) Bioactive Glass Stimulates the Secretion of AngiogenicGrowth Factors and Angiogenesis In Vitro. Tissue Engineering 11:768–777.

[51] Degidi M, Artese L, Scarano A, Perrotti V, Gehrke P, Piattelli A (2006)Inflammatory Infiltrate, Microvessel Density, Nitric Oxide Synthase Expres-sion, Vascular Endotelial Growth Factor Expression, and Proliferative Activityin Peri-implant Soft Tissues around Titanium and Zirconium Oxide HealingCaps. Journal of Periodontology 77:73–80.

[52] Denissen HW, van Dijk HJA, de Groot K, Klopper PJ, Vermeiden JPW, GehringAP (1980) Biological and Mechanical Evaluation of Dense Calcium Hydro-xyapatite Made by Continuous Hot Pressing. in Hastings GW, Williams DF,(Eds.) Mechanical Properties of Biomaterials, Wiley.

[53] Denry I, Kelly R (2008) State of the Art of Zirconia for Dental Applications.Dental Materials 24:299–307.

[54] Doiphode ND, Huang T, Leu MC, Rahaman MN, Day DE (2011) FreezeExtrusion Fabrication of 13–93 Bioactive Glass Scaffolds for Bone Repair.Journal of Materials Science Materials in Medicine 22:515–523.

[55] Domingos M, Chiellini F, Cometa S, Giglio ED, Grillo-Fernandes E, Bartolo PJ,Chiellini E (2010) Evaluation of In Vitro Degradation of PCL Scaffolds Fabri-cated Via BioExtrusion. Part 1: Influence of the Degradation Environment.Virtual and Physical Prototyping 5:1–9.

[56] Domingos M, Chiellini F, Cometa S, De Giglio E, Grilo-Fernandes E, Bartolo PJ,Chiellini E (2011) Evaluation of In Vitro Degradation of PCL Scaffolds Fab-ricated Via BioExtrusion – Part 2: Influence of Pore Size and Geometry. Virtualand Physical Prototyping 6:157–165.

[57] Domingos M, Chiellini F, Gloria A, Ambrosio L, Bartolo PJ, Chiellini E (2012)Effect of Process Parameters on the Morphological and Mechanical Propertiesof 3D Bioextruded Poly(e-caprolactone) Scaffolds. Rapid Prototyping Journal18:56–67.

[58] Domingos M, Dinucci D, Cometa S, Alderighi M, Bartolo PJ, Chiellini F (2009)Polycaprolactone Scaffolds Fabricated Via Bioextrusion for Tissue Engineer-ing Applications. International Journal of Biomaterials 2009:1–9.

[59] Dorozhkin SV (2010) Bioceramics of Calcium Orthophosphates. Biomaterials31:1465–1485.

[60] Ekaputra AK, Zhou Y, Cool SM, Hutmacher DW (2009) Composite ElectrospunScaffolds for Engineering Tubular Bone Grafts. Tissue Engineering Part A15:3779–3788.

[61] Elahinia MH, Hashemi M, Tabesh M, Bhaduri SB (2012) Manufacturing andProcessing of NiTi Implants: A Review. Progress in Materials Science 57(5):911–946.

[62] Elias CN, Lima JHC, Valiev R, Meyers MA (2008) Biomedical Applications ofTitanium and its Alloys. JOM 46–49.

[63] Ferrar B, Mullen L, Jones E, Stamp R, Sutcliffe CJ (2012) Gas Flow Effects onSelective Laser Melting (SLM) Manufacturing Performance. Journal of Materi-als Processing Technology 212:355–364.

[64] Fischbach C, Tessmar J, Lucke A, Schnell E, Schmeer G, Blunk T, Gopferich A(2001) Does UV Radiation Affect Polymer Properties Relevant to TissueEngineering? Surface Science 491:333–345.

[65] Fu Q, Saiz E, Tomsia AP (2011) Direct Ink Writing of Highly Porous and StrongGlass Scaffolds for Load Bearing Bone Defects Repair and Regeneration. ActaBiomaterialia 7:3547–3554.

[66] Gibson I, Rosen DW, Stucker B (2010) Additive Manufacturing Technology:Rapid Prototyping to Direct Digital Manufacturing, Springer.

[72] Gorustovich AA, Roether JA, Boccaccini AR (2010) Effect of Bioactive Glaon Angiogenesis: A Review of In Vitro and In Vivo Evidences. Tissue Engiing Part B Reviews 16:199–207.

[73] Griffith ML, Halloran JW (1996) Freeform Fabrication of Ceramics Via Stelithography. Journal of the American Ceramic Society 79:2601–2608.

[74] Gronet PM, Waskewicz GA, Richardson C (2003) Preformed Acrylic CrImplants Using Fused Deposition Modelling: A Clinical Report. JournProsthetic Dentistry 90(5):429–433.

[75] Hanawa T, Kamiura Y, Yamamoto S, Kohgo T, Anemiya A, Ukai H, MurakaAsoaka K (1997) Early Bone Formation around Calcium-ion Implanted TitanInserted into Rat Tibia. Journal of Biomedical Materials Research 32:131–1

[76] Hao L, Dadbakhsh S, Seaman O, Felstead M (2009) Selective Laser MeltingStainless Steel and Hydroxyapatite Composite for Load-bearing ImpDevelopment. Journal of Materials Processing Technology 209:5793–580

[77] Hao L, Harris R (2008) Customised Implants for Bone ReplacementGrowth. in Bartolo P, Bidanda B, (Eds.) Bio-materials and Prototyping Apptions in Medicine, Springer.

[78] Hao L, Savalani MM, Zhang Y, Tanner KE, Harris RA (2006) Selective LSintering of Nanocomposite Microspheres. Journal of Materials Science Mrials in Medicine 19:2535–2540.

[79] Harrysson OL, Cansizoglu O, Marcellin-Little DJ, Cormier DR, West

(2008) Direct Metal Fabrication of Titanium Implants with Tailored Mateand Mechanical Properties Using Electron Beam Melting Technology. Maals Science and Engineering C 28:366–373.

[80] Hedberg EL, Shih CK, Lemoine JJ, Timmer MD, Liebschner MAK, JanseMikos AG (2005) In Vitro Degradation of Porous Poly(propylene fumarPoly(DL-lactic-co-glycolic acid) Composite Scaffolds. Biomaterials

3215–3225.[81] Hench LL (1991) Bioceramics: From Concept to Clinic. Journal of the Ame

Ceramic Society 74:1487–14510.[82] Hench LL, Wilson J (1993) An Introduction to Bioceramics, World Scien

Publishing.[83] Heymann D, Pradal G, Behahmed M (1999) Cellular Mechanisms of Cal

Phosphate Ceramic Degradation. Histology and Histopathology 14:871–[84] Hollander DA, Walter M, Wirtz T, Sellei R, Schmidt-Rohlfing B, Paar O, Er

(2006) Structural, Mechanical and In Vitro Characterization of IndividStructured Ti–6Al–4V Produced by Direct Laser Forming. Biomate27:955–963.

[85] Hollister SJ, Maddox RD, Taboas JM (2002) Optimal Design and FabricatiScaffolds to Mimic Tissue Properties and Satisfy Biological ConstraBiomaterials 23:4095–4103.

[86] Holy CE, Cheng C, Davies JE, Schoichet MS (2001) Biomaterials 22:25–[87] Hon KKB, Li L, Hutchings IM (2008) Direct Writing Technology – Adva

and Developments. CIRP Annals – Manufacturing Technology 57(2):601–[88] Huang TS, Rahaman MN, Doiphode ND, Leu MC, Bal BS, Day DE, Liu X (2

Porous and Strong Bioactive Glass (13–93) Scaffolds Fabricated by

Extrusion Technique. Materials Science Engineering C 31:1482–1489.[89] Huang W (2002) On the Selection of Shape Memory Alloys for Actua

Materials and Design 23:11–19.[90] Huang YM, Kuriyama S, Jiang CP (2004) Fundamental Study and Theore

Analysis in a Constrained-Surface Stereolitography System. InternatJournal of Advanced Manufacturing Technology 24:361–369.

[91] Huang ZM, Zhangb YZ, Kotakic M, Ramakrishna S (2003) A Review on PolyNanofibers by Electrospinning and their Applications in NanocompoComposites Science and Technology 63:2223–2253.

[92] Hutmacher DW, Schantz JT, Lam CXF, Tan KC, Lim TC (2007) State of thand Future Directions of Scaffold-based Bone Engineering from a Biomials Perspective. Journal of Tissue Engineering and Regenerative Med1:245–260.

[93] Hutmacher DW, Hoque ME, Wong YS (2008) Design, Fabrication and PhyCharacterization of Scaffolds Made from Biodegradable Synthetic PolymeCombination with RP Systems Based on Melt Extrusion. in Bidanda B, BaPJ, (Eds.) Virtual Prototyping & Bio-manufacturing in Medical ApplicatSpringer.

[94] Imwinkelried T (2007) Mechanical Properties of Open-pore Titanium FJournal of Biomedical Materials Research Part A 81:964–970.

[95] Itin VI, Gjunter VE, Shabalovskaya SA, Sachedeva RCL (1994) MechaProperties and Shape Memory of Porous Nitinol. Materials Characteriz32:179.

[96] Jagur-Grodzinski J (2006) Polymers for Tissue Engineering, Medical Devand Regenerative Medicine Concise General Review of Recent Studies.

mers for Advanced Technologies 17:395–418.[97] Jagur-Grodzinski J (2010) Polymeric Gels and Hydrogels for Biomedica

Pharmaceutical Applications. Polymers for Advanced Technology 21:27–[98] Jones A, Vaughan D (2005) Hydrogel Dressings in the Management

Variety of Wound Types: A Review. Journal of Orthopaedic Nursing 9:S1–[99] Jones AC, Arns CH, Hutmacher DW, Milthorpe BK, Sheppard AP, Knack

MA (2009) The Correlation of Pore Morphology, Interconnectivity and

rials

and and

ma-urnal

n thety ofesign

) Theage

[67] Gilbert TW, Sellaro TL, Badylak. (2006) Decellurization of Tissues and Organs.Biomaterials 27:3675–3683.

[68] Gilbert TW, Stewart-Akers AM, Badylak SF (2007) A Quantitative Method forEvaluating the Degradation of Biologic Scaffold Materials. Biomaterials28:147–150.

[69] Goodridge RD, Dalgarno KW, Wood DJ (2006) Indirect Selective Laser Sinter-ing of an Apatite–mullite Glass-ceramic for Potential Use in Bone Replace-ment Applications. Proceedings of the Institution of Mechanical Engineers Part H220:57–68.

[70] Goodridge RD, Wood DJ, Ohtsuki C, Dalgarno KW (2007) Biological Evaluationof an Apatite–mullite Glass-ceramic Produced Via Selective Laser Sintering.Acta Biomaterialia 3:221–231.

[71] Goodridge RD, Tuck CJ, Hague RJM (2012) Laser Sintering of Polyamides andOther Polymers. Progress in Materials Science 57:229–267.


sical Properties of 3D Ceramic Scaffolds with Bone in Growth. Biomate30:1440–1451.

[100] Jurczyk K, Niespodziana K, Jurczyk MU, Jurczyk M (2011) SynthesisCharacterization of Titanium-45S5 Bioglass Nanocomposites. MaterialsDesign 32:2554–2560.

[101] Kamitakahara M, Ohtsuki C, Miyazaki T (2008) Behavior of Ceramic Bioterials Derived from Tricalcium Phosphate in Physiological Condition. Joof Biomaterials Applications 23:197–212.

[102] Kemmoku DT, Noritomi PY, Toland FG, Silva VJVL (2010) Use of BioCAD iDevelopment of a Growth Compliant Prosthetic Device for CranioplasGrowing Patients. in Bartolo PJ, et al. (Eds.) Innovative Developments in Dand Manufacturing, Taylor & Francis.

[103] Kerckhofs G, Pyka G, Loeckx D, Van Bael S, Schrooten J, Wevers M (2010Combined Use of Micro-CT Imaging, In Situ Loading and Non-rigid Im



[104]

[105]

[106]

[107]

[108]

[109]

[110]

[111]

[112]

[113]

[114]

[115]

[116]

[117]

[118]

[119]

[120]

[121]

[122]

[123]

[124]

[125]

[126][127]

[128]

[129]

[130]

[131]

[132]

[133]

[134]


G Model


Plepr

Registration for 3D Experimental Local Strain Mapping on Porous Bone TissueEngineering Scaffolds under Compressive Loading. European Conference forNon-Destructive Testing (ECNDT), Moscow, 7–11.

Kessler-Liechti G, Mericske-Stern R (2006) Rehabilitation of an AbradedOcclusion with Procera-ZrO2 All-ceramic Crowns, A Case Report. SchweizMonatsschr Zahnmed 116:156–167.

Khouja N, Chu T, Chahine G, Kovacevic R, Koike M, Okabe T (2010) In VivoEvaluation of Novel Custom-made Press-fit Implants. International Associa-tion for Dental Research (IADR) General Session, Barcelona, Spain.

Kim H-M, Miyali F, Kobubo T, Nakamura T (1996) Preparation of Ti and itsAlloys Via Simple Chemical Surface Treatment. Journal of Biomedical MaterialsResearch 32:409–417.

Kim JY, Lee JW, Lee AJ, Park EK, Kim SY, Cho DW (2007) Development of aBone Scaffold Using HA Nanopowder and Micro-Stereolithography Technol-ogy. Microelectronic Engineering 84:1762–1765.

Kim SS, Utsunomiya H, Koski JA, Wu BM, Cima MJ, Sohn J, Mukai K, Griffith LG,Vacanti JP (1998) Survival and Function of Hepatocytes o a Novel Three-dimensional Synthetic Biodegradable Polymer Scaffolds with an IntrinsicNetwork of Channels. Annals of Surgery 228:8–13.

Kitsugi T, Nakamura T, Oka M, Yan W-Q, Goto T, Shibuya T, Kobuto T, Miyaji S(1996) Bone Bonding Behaviour of Titanium and its Alloys When Coated withTitanium Oxide (TiO2) and Titanium Silicate (Ti5Si3). Journal of BiomedicalMaterials Research 32:149–156.

Klocke F, Wagner C (2003) Coalescence of Two Metallic Particles as BaseMechanism of Selective Laser Sintering. CIRP Annals – Manufacturing Tech-nology 52(1):177–184.

Koh Y-H, Jun I-K, Kim H-E (2006) Fabrication of Poly(e-caprolactone)/hydro-xyapatite Scaffold Using Rapid Direct Deposition. Materials Letters 60:184–187.

Kohn DH (1998) Metals in Medical Applications. Current Opinion in Solid State& Materials Science 3:309–316.

Koike M, Martinez K, Guo L, Chahine G, Kovacevic R, Okabe T (2011) Evaluationof Titanium Alloy Fabricated Using Electron Beam Melting System for DentalApplications. Journal of Materials Processing Technology 211:1400–1408.

Kokubo T (1992) Bioactivity of Glasses and Glass-ceramics. in Ducheyne P,Kokubo T, van Blitterswijk CA, (Eds.) Bone-bonding Biomaterials, Reed Health-care Communications.

Kolan KCR, Leu MC, Hilmas GE, Brown RF, Velez M (2011) Fabrication of 13–93 Bioactive Glass Scaffolds for Bone Tissue Engineering Using IndirectSelective Laser Sintering. Biofabrication 3:1–5.

Kolan KCR, Leu MC, Hilmas GE, Velez M (2010) Selective Laser Sintering of13–93 Bioactive Glass. Proc. of the 21st Annual Int. Solid Freeform FabricationSymp., Austin, TX, USA, 504–512.

Kolan KCR, Leu MC, Hilmas GE, Velez M (2011) Effect of Particle Size, BinderContent and Heat Treatment on Mechanical Properties of 13–93 BioactiveGlass Scaffolds. Proc. of the 22nd Annual Int. Solid Freeform Fabrication Symp.,Austin, TX, USA, 523–535.

Krishna BV, Bose S, Bandyopadhyay A (2009) Fabrication of Porous NiTi ShapeMemory Alloy Structures Using Laser Engineered Net Shaping. Journal ofBiomedical Materials Research Applied Biomaterials 89:481–490.

Kruth J-P, Vandenbroucke B, Van Vaerenbergh J, Naert I (2005) Rapid Man-ufacturing of Dental Prostheses by Means of Selective Laser Sintering/Melt-ing. Les 11iemes Assises Europeennes du Prototypage Rapide, Paris.

Kruth JP, Levy G, Klocke F, Childs THC (2007) Consolidation Phenomena inLaser and Powder-bed Based Layered Manufacturing. CIRP Annals – Manu-facturing Technology 56(2):730–759.

Kuboki Y, Takita H, Kobayashi D, Tsuruga E, Inoue M, Murata M (1998) BMP-induced Osteogenesis on the Surface of Hydroxyapatite with GeometricallyFeasible and Nonfeasible Structures: Topology of Osteogenesis. Journal ofBiomedical Materials Research 39:190–199.

Kun, X., 2007, Indirect Selective Laser Sintering of Apatite–Wollostonite,Ph.D. Thesis, University of Leeds, UK.

Kyogoku H, Ramos JA, Bourell DL (2002) Laser Melting of Ti–Ni ShapeMemory Alloy. Proc SFF Symp Austin 668–675.

Lam CXF, Teoh SH, Hutmacher DW (2007) Comparison of the Degradation ofPolycaprolactone-(b-tricalcium phosphate) Scaffolds in Alkaline Medium.Polymer International 56:718–728.

Lawrence BJ, Madihally SV (2008) Cell Colonization in 3D Degradable PorousMatrices. Cell Adhesion & Migration 2:9–16.

Learmonth ID (2000) Interface in total hip arthroplasty, Springer. Lee G, Barlow JW (1996) Selective Laser Sintering of Bioceramic Materials for

Implants. 96 SFF Symposium, Austin, TX, August 12–14. Lee J, Cuddihy MJ, Kotov NA (2008) Three-dimensional Cell Culture Matrices:

State of the Art. Tissue Engineering Part B 14:61–86. Lee JJ, Lee SG, Park JC, Yang YI, Kim JK (2007) Investigation of Biodegradabale

PLGA Scaffold with Various Pore Size Structure for Skin Tissue Engineering.Current Applied Physics 7S1:37–40.

Lee JW, Ahn GS, Kim DS, Cho DW (2009) Development of Nano- and Micro-scale Composite 3D Scaffolds Using PPF/DEF-HA and Micro-stereolithogra-

[135] LeGeros RZ (1993) Biodegradation and Bioresorption of Calcium PhosphateCeramics. Clinical Materials 14:65–88.

[136] Leong KF, Chua CK, Sudarmadji N, Yeong WY (2008) Engineering FunctionallyGraded Tissue Engineering Scaffolds. Journal of The Mechanical Behavior ofBiomedical Materials 1:140–152.

[137] Levy RA, Chu TGM, Holloran JW, Feinberg SE, Hollister S (1997) CT-generatedPorous Hydroxyapatite Orbital Floor Prosthesis as a Prototype Bioimplant.American Journal of Neuroradiology 18:1522–1525.

[138] Lewis JA (2000) Colloidal Processing of Ceramics. Journal of the AmericanCeramic Society 83:2341–2359.

[139] Lewis JA, Smay JE, Stuecker J, Cesarano J (2006) Direct Ink Writing of Three-dimensional Ceramic Structures. Langmuir 89:3599–3609.

[140] Lim TW, Park SH, Yang DY (2005) Contour Offset Algorithm for PrecisePatterning in Two-photon Polymerization. Microelectronic Engineering 77:382–388.

[141] Lin FH, Liao CJ, Chen KS, Sun JS, Liu HC (1997) Degradation Behaviour of a NewBioceramic Ca2P2O7 with Addition of Na4P2O7�10H2O. Biomaterials 13:915–921.

[142] Liu CZ, Czernuszka JT (2006) On the Development of Biodegradable Scaffoldsfor Tissue Engineering: A Perspective. Materials Science and Technology12:2479–2488.

[143] Lode A, Wolf-Brandstetter C, Reinstorf A, Bernhardt A, Konig U, Pompe W,Gelinsky M (2007) Calcium Phosphate Bone Cements, Functionalized WithVEGF: Release Kinetics and Biological Activity. Journal of Biomedical MaterialsResearch A 81:474–483.

[144] Lorrison JC, Goodridge RD, Dalgarno KW, Wood DJ (2002) Selective LaserSintering of Bioactive Glass-ceramics. Proc. of the 13th Annual Int. SolidFreeform Fabrication Symp., Austin TX, 1–8.

[145] Lott P, Schleifenbaum H, Meiners W, Wissenbach K, Hinke C, Bultmann J(2011) Design of an Optical System for the In Situ Process Monitoring ofSelective Laser Melting (SLM). Physics Procedia 12:683–690.

[146] Lu L, Garcia CA, Mikos AG (1999) In Vitro Degradation of Thin Poly(DL-lactic-co-glycolic acid) Films. Journal of Biomedical Materials Research 46:244–246.

[147] Machado LG, Savi MA (2003) Medical Applications of Shape Memory Alloys.Brazilian Journal of Medical and Biological Research 36:683–691.

[148] Matsumoto T, Mooney DJ (2006) Cell Instructive Polymers. Advances inBiochemical Engineering/Biotechnology 102:113–137.

[149] Mazzoli A, Germani M, Raffaeli R (2009) Direct Fabrication through ElectronBeam Melting Technology of Custom Cranial Implants Designed in a PHAN-ToM-based Haptic Environment. Materials and Design 30:3186–3192.

[150] Meier H, Zarnetta R, Haberland C, Frenzel J (2009) Selective Laser MeltingNiTi Shape Memory Components. in Bartolo PJ, et al. (Eds.) InnovativeDevelopments in Design and Manufacturing, Taylor & Francis.

[151] Melchels FPW, Domingos MAN, Klein TJ, Malda J, Bartolo PJ, Hutmatcher DW(2012) Additive Manufacturing of Tissues and Organs. Progress in MaterialsScience (http://dx.doi.org/10.1016/j.progpolymsci.2011.11.007, in press).

[152] Mendonca G, Mendonca DBS, Simoes LGP, Araujo AL, Leite ER, Duarte WR,Aragao JL, Cooper LF (2009) The Effects of Implant Surface Nanoscale Featureson Osteoblast-specific Gene Expression. Biomaterials 30:4053–4062.

[153] Meng Q, Hu J (2009) A Review of Shape Memory Polymer Composites andBlends. Composites Part A 40:1661–1672.

[154] Mironov V, Gentile C, Brakke K, Trusk T, Jakab K, Forgacs G, Kasyanov V,Visconti R, Markwald R (2009) Designer Blueprint for Vascular Trees: Mor-phology Evolution of Vascular Tissue Constructs. Virtual and Physical Proto-typing 4:63–74.

[155] Mironov V, Boland T, Trusk T, Forgacs G, Markwald RR (2003) Organ Printing:Computer-aided Jet-based 3D Tissue Engineering. Trends in Biotechnology21:157–161.

[156] Mironov V, Visconti R, Kasyanov V, Forgacs G, Drake C, Markwald R (2009)Organ Printing: Tissue Spheroids as Building Blocks. Biomaterials 30:2164–2174.

[157] Mitchell GR, Ahn KH, Davis FJ (2011) The Potential of Electrospinning inRapid Manufacturing. Virtual and Physical Prototyping 6:63–77.

[158] Moghe AK, Gupta BS (2008) Co-axial Electrospinning for Nanofiber Struc-tures: Preparation and Applications. Polymer Reviews 48:353–377.

[159] Morgan NB (2004) Medical Shape Memory Alloy Applications – The Marketand its Products. Materials Science and Engineering A 378:16–23.

[160] Moroni L, Schotel R, Sohier J, Wijn JR, Blitterswijk CA (2006) Polymer HollowFiber Three-dimensional Matrices With Controllable Cavity and Shell Thick-ness. Biomaterials 27:5918–5926.

[161] Mota C, Puppi D, Dinucci D, Errico C, Bartolo PJ, Chiellini F (2011) Dual-scalePolymeric Constructs as Scaffolds for Tissue Engineering. Materials 4:527–542.

[162] Murr LE, Amato KN, Li SJ, Tian YX, Cheng XY, Gaytan SM, Martinez E, ShindoPW, Medina F, Wicker RB (2011) Microstructure and Mechanical Propertiesof Open-cellular Biomaterials Prototypes for Total Knee ReplacementImplants Fabricated by Electron Beam Melting. Journal of the MechanicalBehavior of Biomedical Materials 4:1396–1411.

[163] Noort RV (2012) The Future of Dental Devices in Digital. Dental Materials
phy. Microelectronics Engineering 86:1465–1467.
Lee JW, Kang KS, Lee SH, Kim JY, Lee BK, Cho DW (2011) Bone RegenerationUsing a Microstereolithography-produced Customized Poly(propylenefumarate)/Dietthyl Fumarate Photopolymer 3D Scaffold IncorporatingBMP-2 Loaded PLGA Microspheres. Biomaterials 32:744–752.

Lee KS, Yang DY, Park SH, Kim RH (2006) Recent Developments in the Use ofTwo-photon Polymerization in Precise 2D and 3D Microfabrication. Polymersfor Advanced Technologies 17:72–82.

Lee KY, Jeong L, Kang YO, Lee SJ, Park WH (2009) Electrospinning of Poly-saccharides for Regenerative Medicine. Advanced Drug Delivery Reviews61:1020–1032.

Lee KY, Park M, Kim HM, Lim YJ, Chun HJ, Kim H, Moon SH (2006) CeramicBioactivity: Progresses, Challenges and Perspectives. Biomedical Materials1:31–37.


28:3–12.[164] O’Brien FJ, Harley BA, Llanas IV, Gibson LJ (2005) The Effect of Pore Size on Cell

Adhesion in Collagen-GAG Scaffolds. Biomaterials 26:433–441.[165] Ovsianikov A, Farsari M, Chichkov BN (2011) Photonic and Biomedical

Applications of the Two-photon Polymerization Technique. in Bartolo PJ,(Ed.) Stereolithographic Processes, Stereolithography: Materials, Processes,Applications, Springer.

[166] Pattanayak DK, Fukuda A, Matsushita T, Takemoto M, Fujibayashi S, Sasaki K,Nishida N, Nakamura T, Kokubo T (2011) Bioactive Ti Metal Analogous toHuman Cancellous Bone: Fabrication by Selective Laser Melting and Chemi-cal Treatments. Acta Biomaterialia 7:1398–1406.

[167] Piperno S, Gheber LA, Canton P, Pich A, Dvorakova G, Biffis A (2009) MicrogelElectrospinning: A Novel Tool for the Fabrication of Nanocomposite Fibers.Polymer 50:6193–6197.


http://dx.doi.org/10.1016/j.progpolymsci.2011.11.007

http://dx.doi.org/10.1016/j.progpolymsci.2011.11.007


008)od-

and

hetic Pro-

fin Surnal

011)ility

andssing

apid008.plexping

andes. vivoaser

Pro-

itionapid

3-E1hate

epla-

nt oftters

ler Rn byrials

nsusork.

bergolac-

26:

ableease.

andissue

signen-

en-hem-

irectdings

ress.eer-

l 11:

andosite

iable

lting

ctur-ce on

tings


G Model


[168] Puleo DA, Kissling RA, Sheu MS (2002) A Technique to Immobilize BioactiveProteins, Including Bone Morphonetic Protein-4 (BMP-4) on Titanium Alloy.Biomaterials 23:2079–2097.

[169] Ragaert K, De Somer F, De Baere I, Cardon L, Degrieck J (2011) Production &Evaluation of PCL Scaffolds for Tissue Engineered Heart Valves. Advances inProduction Engineering & Management Journal 6(3):163–172.

[170] Rainer A, Giannitelli SM, Abbruzzese F, Traversa E, Licoccia S, Trombetta M(2008) Fabrication of Bioactive Glass-ceramic Foams Mimicking Human BonePortions for Regenerative Medicine. Acta Biomaterialia 4:362–369.

[171] Ramakrishna S, Fujihara K, Teo WE, Yong T, Ma Z, Ramaseshan R (2006)Electrospun Nanofibers: Solving Global Issues. Materials Today 9:40–50.

[172] Ramakrishna S, Mayer J, Winterma E (2001) Biomedical Applications ofPolymer Materials: A Review. Composites Science and Technology 61:1189–1224.

[173] Ramsden JJ, Allen DM, Stephenson DJ, Alcock JR, Fuller G, Goch G (2007) TheDesign and Manufacturing of Biomedical Surfaces. CIRP Annals – Manufactur-ing Technology 56(2):687–711.

[174] Ratner BD, Bryant SJ (2004) Biomaterials: Where We Have Been and WhereWe Are Going. Annual Reviews of Biomedical Engineering 6:41–75.

[175] Ratner BD, Hoffman AS, Schoen FJ, Lemons JE (2004) Biomaterials Science: AnIntroduction to Materials in Medicine, 2nd ed. Academic Press.

[176] Rimell JT, Marquis PM (2000) Selective Laser Sintering of Ultra High Mole-cular Weight Polyethylene for Clinical Applications. Journal of BiomedicalMaterials Research 53:414–420.

[177] Rimondini L, Cerroni L, Carrassi A, Torricelli P (2002) Bacterial Colonization ofZirconia Ceramic Surfaces: An In Vitro and In Vivo Study. International Journalof Oral & Maxillofacial Implants 17:793–798.

[178] Ringeisin BR, Othon CM, Barron JA, Young D, Spargo BJ (2006) Jet-basedMethods to Print Living Cells. Biotechnology Journal 1:930–948.

[179] Rouahi M, Champion E, Hardouin P, Anselme K (2006) Quantitative KineticAnalysis of Gene Expression During Human Osteoblastic Adhesion on Ortho-paedic Materials. Biomaterials 27:2829–2844.

[180] Rouwkema J, Rivron NC, van Blitterswijk CA (2008) Vascularization in TissueEngineering. Trends in Biotechnology 26:434–441.

[181] Saito S, Wachi T, Hanada S (1993) A New Fabrication Process of TiNi ShapeMemory Wire. Materials Science and Engineering A 161:91–96.

[182] Sanz-Herrera JA, Garcia-Aznar JM, Doblare M (2009) On Scaffold Designingfor Bone Regeneration: A Computational Multiscale Approach. Acta Bioma-terialia 5:219–229.

[183] Seck T, Melchels FPW, Feijen J, Grijpma DW (2010) Designed BiodegradableHydrogel Structures Prepared by Stereolithography Using Poly(ethylene gly-col)/Poly(D,L-lactide)-based Resins. Journal of Controlled Release 148:34–41.

[184] Shergell H, Kneissi AC (2002) Generation, Development and Degradation ofthe Intrinsic Two-way Shape Memory Effect in Different Alloy Systems. ActaMaterialia 50:327–341.

[185] Shishkovsky I, Kuznettsov MV, Morozov Y (2010) Porous Titanium andNitinol Implants Synthesized by SHS/SLS: Microstuctural and Histomorpho-logical Analysis of Tissue Reactions. International Journal of Self-PropagatingHigh-Temperature Synthesis 19:157–167.

[186] Shishkovsky IV, Volova LT, Kuznetsov MV, Morozovc YG, Parkin IP (2008)Porous Biocompatible Implants and Tissue Scaffolds Synthesized by SelectiveLaser Sintering from Ti and NiTi. Journal of Materials Chemistry 18:1309–1317.

[187] Shu R, McMullen R, Baumann MJ, McCabe LR (2003) Hydroxyapatite Accel-erates Differentiation and Suppresses Growth of MC3T3-E1 Osteoblasts.Journal of Biomedical Materials Research A 67:1196–1204.

[188] Sill TJ, Recum HA (2008) Electrospinning: Applications in Drug Delivery andTissue Engineering. Biomaterials 29(13):1989–2006.

[189] Siow K, Brichter L, Kumar S, Griesser H (2006) Plasma Methods for theGeneration of Chemically Reactive Surfaces for Biomolecule Immobilizationand Cell Colonization. Plasma Processes and Polymers 3:392–418.

[190] Slaughter BV, Khurshid SS, Fischer OZ, Khademhosseini A, Peppas NA (2009)Hydrogels in Regeneration Medicine. Advanced Materials 21:3307–3329.

[191] Smay JE, Cesarano J, Lewis JA (2002) Colloidal Inks for Directed Assembly of 3-D Periodic Structures. Langmuir 18:5429–5437.

[192] Sokolowski W, Metcalfe A, Hayashi S, Yahia L’Hocine. Raymond J (2007)Medical Applications of Shape Memory Polymers. Biomedical Materials2:S23–S27.

[193] Sokolsky-Popkov M, Agashi K, Olaye A, Shakesheff K, Domb AJ (2007) Poly-mer Carriers for Drug Delivery in Tissue Engineering. Advanced Drug DeliveryReviews 59:187–206.

[194] Sun L, Huang WM, Ding Z, Zhao Y, Wang CC, Purnawali H, Tang C (2012)Stimulus-responsive Shape Memory Materials: A Review. Materials andDesign 33:577–640.

[195] Sun L, Wu L, Bao C, Fu C, Wang X, Yao J (2009) Gene Expressions of CollagenType I ALP and BMP-4 in Osteo-inductive BCP Implants Show Similar Patternto that of Natural Healing Bones. Materials Science and Engineering C29:1829–1834.

[200] Tellis BC, Szivek JA, Bliss CL, Margolis DS, Vaidyanathan RK, Calvert P (2Trabecular Scaffold Created Using Micro CT Guided Fused Deposition Melling. Materials Science Engineering C 28:171–178.

[201] Teo WE, Ramakrishna S (2006) A Review on Electrospinning DesignNanofibre Assemblies. Nanotechnology 17:89–106.

[202] Tian H, Tang Z, Zhuang X, Chen X, Jing X (2012) Biodegradable SyntPolymers: Preparation, Functionalization and Biomedical Application.gress in Polymer Science 37:237–280.

[203] Vaccaro AR, Singh K, Haid R, Kitchel S, Wuisaman. Taylor. Branch C, Gar(2003) The Use of Bioabsorbable Implants in the Spine. The Spine Jo3:227–237.

[204] Van Bael S, Kerckhofs G, Moesen M, Pyka G, Schrooten J, Kruth J-P (2Micro-CT-based Improvement of Geometrical and Mechanical Controllabof Selective Laser Melted Ti6Al4V Porous Structures. Materials ScienceEngineering A Structural Materials Properties Microstructure and Proce528(24):7423–7431.

[205] Vandenbroucke, B., Selective Laser Melting of Biocompatible Metals for RManufacturing of Medical Parts, Ph.D. Thesis, K.U. Leuven-PMA, May 2

[206] Vandenbroucke B, Kruth J-P (2008) Direct Digital Manufacturing of ComDental Prostheses. in Bartolo P, Bidanda B, (Eds.) Bio-materials and PrototyApplications in Medicine, Springer.

[207] Vandenbroucke, B., Kruth, J.-P., Naert, I., 2005, Procedure for DesignProduction of Implant-based Framework for Complex Dental Prosthes

[208] Velez M, Jung S, Kolan KCR, Leu MC, Day DE, Chu T-MG (2011) InEvaluation of 13–93 Bioactive Glass Scaffolds Made by Selective LSintering. Next Generation Biomaterials Symposium, Ceramic Transactionsceedings, Materials Science and Technology, Columbus, OH, USA.

[209] Wang F, Shor L, Darling A, Khalil S, Guceri S, Lau A (2004) Precision Deposand Characterisation of Cellular Poly-e-caprolactone Tissue Scaffolds. RPrototyping Journal 10:42–49.

[210] Wang J (2009) Proliferation and Differentiation of Osteoblast-like MC3TCells on Biomimetically and Electrolytically Deposited Calcium PhospCoatings. Journal of Biomedical Materials Research A 90:664–670.

[211] Wang M (2003) Developing Bioactive Composite Materials for Tissue Rcement. Biomaterials 24:2133–2151.

[212] Wang Y, Shen Y, Wang Z, Yang J, Liu N, Huang W (2010) DevelopmeHighly Porous Titanium Scaffolds by Selective Laser Melting. Materials Le64:674–675.

[213] Wenisch S, Stahl JP, Horas U, Heiss C, Kilian O, Trinkaus K, Hild A, Schnett(2003) In Vivo Mechanisms of Hydroxyapatite Ceramic DegradatioOsteoclasts: Fine Structural Microscopy. Journal of Biomedical MateResearch Part A 67A:713–718.

[214] Williams DF (1987) Definitions in Biomaterials. Proceedings of a ConseConference of the European Society for Biomaterials, vol. 4Elsevier, New Y

[215] Williams JM, Adewunmi A, Schek RM, Flanagan CL, Krebsbach PH, FeinSE, Hollister SJ, Das S (2005) Bone Tissue Engineering Using Polycaprtone Scaffolds Fabricated Via Selective Laser Sintering. Biomaterials4817–4827.

[216] Wischke A, Neffe AT, Steuer S, Lendlein A (2009) Evaluation of a DegradShape-memory Polymer Network as Matrix for Controlled Drug RelJournal of Controlled Release 138:242–250.

[217] Wood DJ, Dyson J, Xiao K, Dalgarno KW, Genever P (2008) ProcessingCharacterization of Apatite–Wollostonite Porous Scaffolds for Bone TEngineering. Key Engineering Materials 361–363:923–926.

[218] Woodfield TB, Malda J, Wijn J, Peters F, Riesle J, Blitterswijk CA (2004) Deof Porous Scaffolds for Cartilage Tissue Engineering Using a Three-dimsional Fiber-deposition. Biomaterials 25:4149–4161.

[219] Wu S, Serbin J, Gu M (2005) Two-photon Polymerization for Three-dimsional Micro-fabrication. Journal of Photochemistry and Photobiology A Cistry 181:1–11.

[220] Xiao K, Dalgarno KW, Wood DJ, Goodridge RD, Ohtsuki C (2008) IndSelective Laser Sintering of Apatite–wollostonite Glass-ceramic. Proceeof the Institution of Mechanical Engineers Part H 222:1107–1114.

[221] Xie, T., 2011, Recent Advances in Polymer Shape Memory, Polymer, in p[222] Xiong Z, Yan Y, Zhang R, Wang X (2005) Organism Manufacturing Engin

ing Based on Rapid Prototyping Principles. Rapid Prototyping Journa160–166.

[223] Xu C, Su P, Chen X, Meng Y, Yu W, Xiang AP, Wang Y (2011) BiocompatibilityOsteogenesis of Biomimetic Bioglass–Collagen–Phosphatidylserine CompScaffolds for Bone Tissue Engineering. Biomaterials 32:1051–1058.

[224] Xu, T., Jin, J., Gregory, C., Hickman, J.J., Boland, T., Inkjet Printing of VMammalian Cells, Biomaterials, 26:93–99.

[225] Yadroitsev I, Smurov I (2011) Surface Morphology in Selective Laser Meof Metal Powders. Physics Procedia 12:264–270.

[226] Yan Y, Zhang R, Lin F (2003) Research and Applications on Bio-manufaing. in Bartolo PJ, et al. (Eds.) Proceedings of the 1st International ConferenAdvanced Research in Virtual and Rapid Prototyping, .

[227] Yang Y, Kim KH, Ong JL (2005) A Review on Calcium Phosphate Coa
ying.
ancetistry

ts ofatrix

aserrbo-rials

[196] Sun L, Berndt CC, Gross KA, Kucuk A (2001) Review: Material Fundamentalsand Clinical Performance of Plasma Sprayed Hydroxyapatite Coatings. Journalof Biomedical Materials Research 58:570–592.

[197] Sung HJ, Meredith C, Johnson C, Galis ZS (2004) The Effect of ScaffoldDegradation Rate on Three-dimensional Cell Growth and Angiogenesis.Biomaterials 25:5735–5742.

[198] Suska F, Kalltorp M, Esposito M, Gretzer C, Tengvall P, Thomsen P (2001) Invivo/Ex Vivo Cellular Interactions With Titanium and Copper. Journal ofMaterials Science Materials in Medicine 12:939–944.

[199] Tan KH, Chua CK, Leong KF, Cheah CM, Cheang P, Abu Bakar MS, Cha SW (2003)Scaffold Development Using Selective Laser Sintering of Polyetheretherketone-Hydroxyapatite Biocomposite Blends. Biomaterials 24:3115–3123.


Produced Using a Sputtering Process – An Alternative to Plasma SpraBiomaterials 26:327–337.

[228] Yildirim M, Fischer H, Marx R, Edelhoff D (2003) In Vivo Fracture Resistof Implant Supported All-ceramic Restorations. Journal of Prosthetic Den90:325–331.

[229] Zeltinger J, Sheerwood JK, Graham DM, Mueller R, Griffith LG (2001) EffecPore Size and Void Fraction on Cellular Adhesion, Proliferation and MDeposition. Tissue Engineering 7:557–572.

[230] Zhou WY, Lee SH, Wang M, Cheung WL, Ip WY (2008) Selective LSintering of Porous Tissue Engineering Scaffolds from Poly(L-lactide)/Canated Hydroxyapatite Nanocomposite Microspheres. Journal of MateScience Materials in Medicine 19:2535–2540.


http://dx.doi.org/10.1016/j.polymer.2011.08.003

http://dx.doi.org/10.1016/j.polymer.2011.08.003


Date post:	01-Jan-2017
Category:	Documents
Upload:	lamanh
View:	233 times
Download:	8 times

Biomedical production of implants by additive electro-chemical and ...

Documents