biodegradable materials for clinical applications: a...

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Reviews in Advanced Sciences and EngineeringVol. 4, pp. 1–18, 2015

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Biodegradable Materials for ClinicalApplications: A ReviewMuralidharan Paramsothy1,∗ and Seeram Ramakrishna1,2

1Centre for Nanofibers and Nanotechnology, Department of Mechanical Engineering,National University of Singapore, 9 Engineering Drive 1, 117576, Singapore2Institute for CNS Regeneration, Jinan University, Guangzhou, 510632, China

ABSTRACT

This review article focuses on mechanical properties and composition of degradable polymer, metal (magne-sium), ceramic and composite materials with respect to clinical application involving cortical/cancellous bone,dentin and enamel, and ligament, tendon and fascia. On the basis of mechanical property comparison, eachclass of densified material may be used to substitute cortical bone, may not be used to substitute cancel-lous bone, may be used to substitute enamel and dentin, may be used to substitute lower limb ligaments andtendons, and may be used to substitute selected upper limb and trunk ligaments and associated tissues. Forrelatively longer times of recovery from more severe orthopedic trauma, PLLA (poly L-lactic acid) and TMC(trimethyl carbonate) polymers, selected Mg alloys containing zinc (Zn), calcium (Ca) and/or rare-earth (RE)elements, or hydroxyapatite (HA) and tricalcium phosphate (TCP) as crystalline bioceramics apparently aremost suitable. Mg alloys or composites have significant potential for clinical application where the ability to bearappropriate load at least in the initial stages of recovery prior to significant resorption and load transfer to newtissue is critical. However, there is no clear opinion at present regarding the toxicity levels of many alloyingelements in Mg alloys, particularly concerning RE elements. There is also the potential issue of nanoparticlecytotoxicity concerning alloy nanocomposite degradation invivo. Compared to crystalline bioceramics, the higherindex of bioactivity (IB) of amorphous glasses and glass ceramics reflects their superior ability to form a denseinterphase and bond strongly with bone, despite their mechanical inferiority. The combination of bioceramicand biodegradable polymer is synergistic based on direct improvement of mechanical property and degradationresistance of the polymer, indirect reduction of foreign body interactions, and direct increase in toughness ofthe ceramic. Future directions of this field regarding developments on cellulose as a biodegradable material,bone tissue regeneration and engineering, stronger and more corrosion resistant and biocompatible magne-sium alloy systems, applicable nanoparticles and nanotechnology, and soft tissue engineering such as vasculartissue engineering are also addressed.

KEYWORDS: Biodegradable, Bone, Enamel, Dentin, Collagen, Magnesium, Bioceramic, Nanoparticles.

CONTENTS1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12. Basic Mechanical Properties of Cortical/Cancellous Bone,

Dentin and Enamel, and Ligament, Tendon and Fascia . . . . . . 33. Biodegradable Polymers . . . . . . . . . . . . . . . . . . . . . . . . . 34. Biodegradable Metals (Magnesium Alloys) . . . . . . . . . . . . . 65. Biodegradable Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . 106. Biodegradable Composites . . . . . . . . . . . . . . . . . . . . . . . . 127. Future Directions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 138. Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 14

References and Notes . . . . . . . . . . . . . . . . . . . . . . . . . . . 15

∗Author to whom correspondence should be addressed.Email: [email protected]: xx Xxxx xxxxAccepted: xx Xxxx xxxx

1. INTRODUCTIONBiomaterials are used to make devices to replace a part ora function of the body in a safe, reliably economical, andphysiologically acceptable manner. A variety of devicesand materials are used in the treatment of disease or injury.Some of these devices remain in the body without signif-icant degradation for extended lengths of time, e.g., tita-nium alloy based pacemaker and arthroplastic devices forthe hip and knee joints.1�2 There are also devices sur-gically removed after planned durations of implantationwhen the tissue has healed, e.g., various bone fixationplates and screws made of stainless steel.3�4 Further, somedevices may be left in the body to intentionally degradeduring tissue healing, e.g., polymeric implants for menis-cus repair, selected metallic (magnesium based) orthopedic

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Biodegradable Materials for Clinical Applications: A Review Paramsothy and Ramakrishna

screws and phosphate ceramic based bone inserts.5�6 Thislast group of devices is of significant appeal based on theelimination of the surgical removal step (and significantreduction of surgical risk) towards complete healing.Devices made of biodegradable materials are designed

based on mechanical properties as well as composi-tion. Relevant parts of the body that biodegradablematerials are applicable for clinical application includecortical/cancellous bone, dentin and enamel, cartilage, andligament, tendon and fascia.7�8 Generally, it is importantto match the stiffness of the targeted body part to avoidharmful effects like stress shielding.9�10 In stress shield-ing due to much higher stiffness of the implanted devicecompared to the surrounding tissue, areas adjacent to theimplanted biodegradable device tend to heal well but atthe expense of regions further away from the implanteddevice. The higher stiffness of the implanted device pointstoward adjacent areas preferentially conducting higher

Muralidharan Paramsothy is currently Associate Faculty of the School of Scienceand Technology at the Singapore Institute of Management University (UNISIM) wherehe teaches topics involving biomedical engineering and materials science. He is cur-rently also Scientist/Consultant at NanoWorld Innovations (NWI): http://www.nanoparticle-magic.vpweb.com. He obtained his Ph.D. in the Department of Mechanical Engineeringat the National University of Singapore in early 2010 and has over 3 years of postdoc-toral experience. With about 60 journal papers and about 20 conference papers (leading to∼800 citations, H-Index of 13, and i-10-Index of 21), his research is focused on the ex-situand in-situ addition of nanoparticles to magnesium and aluminium alloys as well as selectedpolymer and bioceramic materials, including nanoparticle-matrix interactions at nanoscalewhich simultaneously give rise to beneficial material properties at macroscale. He is the

founder of the Nanoscale Electro Negative Interface Density or NENID theory applicable to abundant interface nanoscalesystems, is on the Editorial Boards of Surface Engineering (Maney Publishing, UK), Nanoscience and Technology: OpenAccess (Symbiosis Group, USA), Journal of Research in Nanotechnology (IBIMA Publishing, USA), and is an activemember of the Composite Materials Committee of The Minerals, Metals and Materials Society (TMS).

Seeram Ramakrishna PE, FREng is well known to the worldwide community of electro-spinners and materials. He advances science and engineering of nanomaterials for innova-tions in areas of societal importance. He is a professor of materials engineering and Directorof Center for Nanofibers and Nanotechnology at the National University of Singapore.He is a Highly Cited Materials Scientist. He authored 11 books and ∼1000 journalpapers which attracted ∼44,000 citations and 98 H-index (with 470 i10-Index). He is onthe editorial boards of ∼10 international journals. Companies Biomers (http://simpliclear.com/), Electrospunra (http://www.electrospunra.com/#page_2/) and Electrospin Tech(http://electrospintech.com/) have roots in the technologies nurtured at his labs. He receivedacademic training from Cambridge and Harvard Universities. He is an elected internationalfellow of Royal Academy of Engineering, UK; National Academy of Engineering, India;

Institution of Engineers Singapore; ASEAN Academy of Engineering and Technology; American Association of theAdvancement of Science; ASM International; American Society for Mechanical Engineers; American Institute for Medi-cal and Biological Engineering; Institution of Mechanical Engineers, UK; and Institute of Materials, Minerals and Mining,UK. He is an analyst and speaker at the meetings facilitated by UNESCO, World Bank, OECD, EU, ASEAN, Gov-ernments, Universities, and professional organizations around the world, with global contributions such as Founder ofGlobal Engineering Deans Council (www.gedc.org); Vice-President of International Federation of Engineering EducationSocieties (IFEES); Board Member of Asia Society for Innovation and Policy (ASIP); and Council Member of Institutionof Engineers Singapore (IES). His university leadership includes National University of Singapore (NUS) Vice-Presidentfor Research Strategy; Dean of Faculty of Engineering.

levels of mechanical stress compared to regions furtheraway from the implanted device. The degradation of poly-mer, metal and ceramic under physiological conditions isalso different for each class of material. Polymers tendto degrade invivo due to ionic attack, oxygen dissolu-tion, enzyme attack and hydration, implying the lower sta-bility of hydrophilic polymers compared to hydrophilicpolymers.11�12 Chain scission in the polymer backbonemechanizes the degradation. Polymer degradation effectsinvivo include undesirable tissue reactions when the degra-dation rates are high. Metals degrade invivo primarilyaccording to corrosion mechanisms involving metal ionrelease.13�14 Excessive corrosion rates contribute to highmetal ion uptake by tissue adjacent to the metal implantdevice. The high metal ion uptake generally leads to unde-sirable adjacent tissue reactions. Ceramics degrade invivoby leaching of selected network forming elements.15�16 Theleaching also involves ion formation from the elements and

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Paramsothy and Ramakrishna Biodegradable Materials for Clinical Applications: A Review

excessive ion uptake also similarly leads to undesirableadjacent tissue reactions.

The key to successfully designing invivo biodegradableimplant devices involves (in part):(1) tailoring of mechanical properties and(2) effective regulation of degradation of the material(s)used in the implant device.

It is also necessary to understand the physiology of thebody as well as patho-physiological response of the bodyto various elements, linked to short term as well aslong term effects of degradable device implantation.17�18

This review article focuses on mechanical properties andcomposition of degradable polymer, metal (magnesium),ceramic and composite materials with respect to clinicalapplication involving cortical/cancellous bone, dentin andenamel, and ligament, tendon and fascia.

2. BASIC MECHANICAL PROPERTIES OFCORTICAL/CANCELLOUS BONE, DENTINAND ENAMEL, AND LIGAMENT, TENDONAND FASCIA

In cortical bone, secondary osteons are roughly cylindri-cal structures that are typically several millimeters longand around 0.2 mm in diameter. There are longitudinalosteons and alternate osteons whose collagen fibers arelongitudinally oriented and non-longitudinally (i.e., alter-nately) oriented, respectively.19–21 Table I lists the elasticmoduli and strengths of longitudinal osteons and alternateosteons. Generally, bending modulus is lowest while tor-sional modulus is highest, and bending strength is over-all highest compared to strengths in other deformationmodes.

In cancellous bone, also known as trabecular or spongybone, there are lamellar bone packets of rod and platearrays that make up interconnected passages which arefilled with bone marrow.21 Table II lists the moduliand compressive strengths of cancellous bone. Hav-ing a heterogeneous open cell porous structure, can-cellous bone exhibits anisotropic mechanical propertiesthat depend on porosity and architectural arrangement of

Table I. Elastic moduli and strengths of longitudinal and alternateosteons.19–21

Longitudinal osteons Alternate osteons

Modulus (GPa)Tension 11�7 5�5Compression 6�3 7�4Bending 2�3 2�6Torsional 22�7 16�8

Strength (MPa)Tension 120 102Compression 110 134Bending 390 348Torsional 202 167

Table II. Modulus and compressive strengths of cancellous bone.21� 22

Tissue source Modulus (GPa) Ultimate strength (MPa)

Proximal Tibia 0.445 (0.257) 5.33 (2.93)Femur 0.389 (0.270) 7.36 (4.00)Lumbar Spine 0.067 (0.044) 2.45 (1.52)

0.023 (0.015) 1.55 (1.11)

Note: Standard deviations are given in parentheses.

the interconnected passages. Modulus is in the range of0.010–2 GPa while strength is in the range of 0.1–30 MPa,where strength is linearly and strongly correlated withmodulus.21�22

Enamel and dentin pertain to teeth, i.e., incisors,canines, premolars and molars, which are used for cutting,tearing, grasping and grinding food, respectively. Teethare structurally heterogeneous where harder and strongerenamel coat softer yet tougher dentin, where there isthe dentino-enamel-juction (DEJ) as the interphase.21�23

Table III lists the basic mechanical properties of enameland dentin. Generally, modulus and strength of enamel inmolars is highest, and modulus and strength of dentin iscomparatively invariant.Ligament, tendon and fascia consist primarily of colla-

gen fibers. Ligament connects bone to bone, tendon con-nects muscle to bone, and fascia is a sheet of fibroustissue which encloses muscle. In ligaments and tendons,the collagen fibers are in aligned bundle packs whereas infascia, there may be no dominant alignment of collagenfibers.21�24�25 Tables IV and V list the basic mechanicalproperties of lower limb ligaments, tendons and fascia,and upper limb and trunk ligaments and associated tis-sues, respectively. Generally, the ligaments and associatedtissues in the lower limbs tend to have higher modulus,higher strength and lower strain at ultimate strength, com-pared to the ligaments and associated tissues in the upperlimbs.

3. BIODEGRADABLE POLYMERSThe main biodegradable polymers applicable to clini-cal applications are poly(lactic acid) PLA, poly(glcolicacid) PGA and PLA/PGA copolymers.36–46 Other poly-mers include collagen, poly (hydroxybutyrate) PHB, poly-carbonates derived from desaminotyrosyl-tyrosine ethylester DTE, cellulose acetate CA, polydioxanone PDS,polycaprolactone PCL, polyphosphazines, chitin or choto-san, poly(ethylene glycol) PEG, poly(vinyl alcohol) PVA,polyethylene oxide-polybutylene terephthalate PEO/PBTcopolymer, and polypropylene fumarate (PPF).47–57

Collagen is a major component of tissue and accountsfor 30% of protein in the body. Every tissue requiringstrength and flexibility has collagen and there are manyforms of collagen in the body. Generally, collagen con-sists of triple helix formed proteins spanned across asignificant portion of the molecule. The molecules may

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easily polymerise into fibers or larger structures and mayalso be further modified with intra/inter-molecular links toform the macroscopic fibers, fibrils, and bundles to formtissue. Disadvantageously, cross linking generally leadsto lower deformability and increases tendency towardscalcification.50

Chitosan is the deacylated derivative of chitin, which isa naturally occurring polysaccharide. The residual acetylcontent determines the biodegradation rate of chitosanand can be easily tailored during synthesis, enablingeasy synthesis of chitosan with varied degradation rates.The chitosan macrostructure is sensitive to pH. It forms agel above a pH of 6 in the presence of phosphate ions,and may form polyelectrolyte complexes with proteinsand polysaccharides for use in controlled drug deliveryapplications. Other bioapplications include encapsulation,membrane barriers, contact lens material, and inhibitorsof blood coagulation as opposed to heparin.51 Chemicallymodified chitosan is a well tolerated implant material espe-cially in tissue regenerative therapy.Polyhdroxybutyrate (PHB) is a very simple form of

polyhydroxyalkanoate (PHA). PHB is a storage polymerin organisms where its role is to store carbon dioxide andenergy. PHB is highly crystalline, very brittle, hydropho-bic, and easily undergoes hydrolytic or enzymatic degra-dation. However, PHB degradation invivo can take years.When PHB contains hydroxyvaleric acid, it is less crys-talline, more flexible, and therefore may be more easilyprocessed. PHB degrades invivo to a natural componentof blood so is considered to have low toxicity. PHB hasbeen used in controlled drug release applications, includ-ing sutures and artificial skin.48

Polylactic acid (PLA), polyglycolic acid (PGA), andderived copolymers have been extensively used as safesynthetic biodegradable polymers in a variety of clini-cal applications. PLA, PGA and derived copolymers arethe benchmarked biopolymers in the biomedical indus-try. PGA is highly crystalline, has a high melting point,and has low solubility in organic solvents. It is com-mercially available under the name Dexon, and was firstimplemented as biodegradable sutures 4 decades ago.PLA is more hydrophobic than PGA and was copoly-merized with it for more biomedical industrial applica-bility. Copolymer sutures include Vicryl and Polyglactin910. Interstingly, the high hydrophobic nature of PLAand highly crystalline nature of PGA are both lost whencombined to form copolymers. PGA/PLA copolymersgenerally have lower crystallinity and hydrophobicity com-pared to the individual polymers, regardless of mixingratio, and degrade during clinical use more rapidly than theindividual polymers. Being esters, the degradation causeslocalized pH drop which may auto-accelerate the degrada-tion and induce inflammatory response.58–60 However, evenwhen the degradation is controlled to avoid this disad-vantage, another disadvantage is the non-robust but rather

weak nature of porous structure arising from PGA/PLAcopolymerization which limits use in hard tissue regener-ation. There has been focus on PLA and PLA/PGA platesand screws for bone fixation, taking these disadvantagesinto account.36�37�43�48 PLA/PGA based polymers may beprocessed by standard injection and compression moldingbut melt crystallization during processing leads to infe-rior mechanical strength.32�36 Self-reinforcing (SR) tech-niques have been used, namely sintering and fibrillation,to increase mechanical strength leading to better clini-cal applicability.61�62 SR processed PLA and PLA/PGAsystems have found application as orthopedic plates andscrews under the Biofix tradename.63

Polycaprolactone (PCL) is a semicrystalline polymerwith a low melting point and low glass transition tempera-ture. PCL also has high thermal stability and easily formsblends with a wide range of polymers. PCL degrades ata slower rate compared to PLA and can be used in drugdelivery devices active for a year. PCL is considered non-toxic and tissue compatible, and is often used as a tissuescaffold.52–54

Polydioxanone (PDS) has poor mechanical strength,55% crystallinity, and a glass transition temperatureslightly below 0 �C. It cannot be used for manufactur-ing medical implants. PDS does not exhibit acute toxiceffects, but resorbs within 6 months, making it compatibleas dressing for slow-healing wounds.48

Polyethylene oxide-polybutylene terephthalate PEO/PBTcopolymer have been investigated for bone replace-ment applications. Weighting of each component allowscontrolled determination of mechanical properties anddegradation characteristics towards good bone bondingbehavior based on favorable hydrogel properties. For softtissue applications like skin scaffold, suitable mechani-cal properties and good biocompatibility invitro and invivohave been demonstrated.64

Polypropylene fumarate (PPF) is viscous at room tem-perature but may be cross-linked with N-VinylPyrolidone(N-VP) releasing low amounts of heat to form putty priorto injection or molding into the bone defect of com-plex geometry. This is considered near net shape pro-cessing during surgery where compared with other formsof polymerization, the heat released during putty forma-tion and injection/molding is significantly lower, there-fore reducing the risk of local tissue necrosis as observedinvivo.55

Table VI lists the mechanical properties and degra-dation data of biodegradable polymers. Compared withTables I–V, based on basic mechanical properties, mostpolymers in Table VI may be used to substitute corticalbone, may not be used to substitute cancellous bone, maybe used to substitute enamel and dentin, may be used tosubstitute lower limb ligaments and tendons, and may beused to substitute selected upper limb and trunk ligamentsand associated tissues. Table VI also lists the commer-cially available resorbable or bioabsorbable devices for

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Table III. Elastic modulus and strengths of enamel and dentin.21� 23

Incisor Canine Pre-molar Molar

Modulus (GPa)Enamel 84�3 (8.9)

77�9 (4.8)48 (6) 46 (5)33 (2) 32 (4)

9�7 (3)12 (3)

Dentin 11�0 (5.8)13 (4) 14 (6) 14 (0.7) 12 (2)9.7 (2) 12 (3) 9.0 (2) 7�6 (3)

10.1610.879.49

Strength (MPa)Enamel, stress at 353 (83)proportional limit

336 (61)194 (19) 224 (26)183 (12) 186�2 (17)

70�3 (22)98�6 (26)

91.0 (10)Enamel, tensile 10 (2.6)strength

Enamel, compressive 384 (92)strength

372 (56)288 (48) 261 (41)253 (35) 239 (30)

94�5 (32)127 (30)

220 (13)

Dentin, stress at 167 (20.0)proportional limit

124 (26) 140 (15) 146 (17) 148 (21)86 (24) 112 (34) 110 (38) 108 (39)

110�5 (22.6)167�3 (37.5)103�1 (16.8)158 (32)154 (23)

Dentin, tensile 52 (10)strength

37�3 (13.6)34�5 (11.1)37�3 (9.0)39�3 (7.4)

Dentin, compressive 297 (24.8)strength

232 (21) 276 (72) 248 (10) 305 (59)233 (66) 217 (26) 231 (38) 250 (60)

295 (21)251 (30)

Note: Standard deviations are given in parentheses.

osteofixation. This reinforces the logic behind which typeof polymer to use for clinical osteo-applications, depend-ing on intended duration of patient recovery. Apparently,PLLA (poly L-lactic acid) and TMC (trimethyl carbonate)

Table IV. Mechanical properties of lower limb ligaments, tendons andfascia.21� 24� 25

Modulus Ultimate tensile Strain at ultimateTissue (MPa) strength (MPa) tensile strength (%)

KneeLigamentAnt. cruciate 65–541 13–46 9–44Pos. cruc 109–413 24–36 10–29

TendonPatellar 143–660 24–69 14–27

AnkleLigamentLat. collateral 216–512 24–46 13–17Med. collateral 54–321 16–34 10–33

TendonAchilles 65 24–61 24–59Palmaris long. 2310±620 91±15 NA

OtherTendonSemitend. 362±22 89±5 52±3Gracilis 613±41 112±4 34±2

FasciaTibial 283±132 14±4 NAFascia lata 150–571 30–105 27–29

Notes: Ant for anterior, Pos for posterior, Med for medial, Lat for lateral, Long forlongus, Semitend for semitendenosis, NA for not available.

are best suited for longer durations of patient recoverywhile PGA (poly glycolic acid) and PDS (poly dioxanone)are best suited for relatively shorter durations of patientrecovery.

Table V. Mechanical properties of upper limb and trunk ligaments andassociated tissues.21� 26� 27

Modulus Ultimate Tensile Strain at ultimateTissue (MPa) Strength (MPa) tensile strength (%)

ShoulderLigamentInf. glenohum. 30–42 5–6 8–15Capsule 32–67 8–21 NA

SpineLigamentPos. long. NA 21–28 11–44Liga flavum NA 1–15 21–102Ant. long. 286–724 8–37 10–57Supraspinal NA 9–16 39–115Interspinal NA 2–9 39–120Intertransverse NA 51±1.4 16.5±0.7

ForearmLigamentCarpal joint 23–119 NA NAPalmar radioul. 39±18 5.7±1.7 51±24Dorsal radioul. 52±33 8±5 61±29

InterosseousMembrane 528±82 43±1.4 10±2

Notes: Inf for inferior, glenohum. for glenohumoral, Pos for posterior, long. forlongus, Ant for anterior, radioul for radioulnar, NA for not available.

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Table VI. Mechanical properties and degradation data of biodegradable polymers, including commercially available resorbable or bioabsorbabledevices for osteofixation.28–35� 63

Tensile Tensile Bending MechanicalPolymer modulus (GPa) strength (MPa) strength (MPa) Weight loss property loss

PGA 6.5–7.0 57–100 120–218 100% in 60–80 days –PLA or PLLA 2.0–5.1 10–75 45–145 100% in 2–5 years 50% in 1 month for

semi-crystalline and in 5months for amorphous

PDLLA 3.5 60 130 100% in 0.5–1 year –Sintered self-reinforced

PGA (2–4.5 mm dia.)7–10 NA 200–260 – 67% in 3 weeks. Loss below

cancellous bone levels in 4–7weeks.

Hot drawn self-reinforcedPGA (2–4.5 mm dia.)

13 NA 360 – 67% in 3 weeks. Loss belowcancellous bone levels in 4–7weeks.

Self-reinforced PLLA 7–8 120 245–275 – 60% in 12 weeksHot drawn PLLA – – – – 70% in 24 weeksPHB 4 40 NA NA NAPolycarbonates (DTE) 1.4–1.8 40–60 NA 40% in 16 weeks. 50%

in 1 year.–

CA 1.6 26.5 NA 6% in 180 days –PDS 2 30 NA 100% in 180 days –

Commercially available resorbable or bioabsorbable devices for osteofixation

Product Manufacturer Year Conformation Biodegradation period

Biofix® BionX 1984 SR-PGA 6 weeksOrthosorb® DePuy 1991 PDS 6 monthsFixsorbMX® Takiron 1994 PLLA 2 yearsLactosorb® Walter Lorenz 1996 PLLA/PGA 12–18 monthsMacroSorb® Macropore 1999 P-L/D-LA 2 yearsResorbX® KLS martin 2001 P-L/D-LA 2 yearsInion CPS® Inion 2001 P-L/D-LA 2 yearsBiosorbFX® Bionix Implants 2001 P-L/D-LA 2 yearsPolyMax® Synthes 2003 P-L/D-LA 2 yearsDelta System® Stryker 2004 P-L/D-LA /GA 2 yearsOsteotransMX® Takiron 2007 u-HA/PLLA 5.5 yearsInion CPS® Inion 2007 P-L/D-LA/TMC 2 years

Notes: PGA: Poly(Glycolic Acid), PLA: Poly(Lactic Acid), PLLA: Poly(L-Lactic Acid), PDLLA: Poly(D,L-Lactic Acid), PHB: Poly(HydroxyButyrate), DTE:Desaminotyrosyl-Tyrosine Ethyl Ester, CA: Cellulose Acetate, PDS, PolyDioxanone, NA: Not Available, SR: Self Reinforced, u-HA: unsintered HydroxyApatite, TMC:TriMethylCarbonate.

4. BIODEGRADABLE METALS(MAGNESIUM ALLOYS)

Magnesium (Mg) is the central atom in the photosynthesisprocess occurring in all plants under sunlight containingthe green pigment chlorophyll, making it one of the mainbuilding blocks of nature next to carbon and silicon. Inparallel fashion, as the lightest structural metal known toman, magnesium has valuable application in the aerospace,automotive, consumer electronics, sports and recreational,and very importantly defence industries.65 In comparison,recent studies have cited Mg and its alloys as potentialmaterials for biomedical applications.66–74 The advantagesMg-based materials possess include apparent non-toxicity,biocompatibility, mechanical compatibility with bone, andbiodegradability in the body.75

Concerning good biological behaviour, Mg is an essen-tial element for the human body. Being the fourth most

abundant cation in the human body, Mg2+ is stored mainlyin bone tissue. Mg2+ is the direct corrosion product ofmagnesium, is absorbed by the human body easily, andcan be excreted in urine.75 Concerning good mechan-ical properties, Mg based metals have advantage overthe currently developed biodegradable materials such aspolymers and ceramics (including bioactive glasses) inload bearing applications requiring higher strengths. Mgalso has Young’s modulus similar or close to that ofbone (in the 20–40 GPa range), compared with tita-nium (Ti) alloys (about 110 GPa), stainless steel (SS)alloys (about 200 GPa) and cobalt (Co) based alloys (about230 GPa). This generally mitigates or minimizes the stressshielding effect induced by serious mismatch in modu-lus between natural bone and metal implant. The den-sity of Mg (1.74 g/cc) is also close to those of naturalbone (1.8–2.1 g/cc), compared with Ti alloys (4.42 g/cc

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for Ti–6Al-4V), SS alloys (about 7.8 g/cc), biodegradablepolymers (about 1 g/cc for PLLA) and hydroxyapatite(3.16 g/cc).

Regarding metallurgical categories of Mg alloys,ZK60A (Mg–6Zn–0.5Zr) is commonly used in structuralapplications based on:(a) high strength and ductility after T5 aging,(b) good creep resistance,(c) poor arc weldability due to hot-shortness cracking and(d) excellent resistance weldability.65

Overall, Mg–Zn alloys are well known for their precip-itation hardening characteristics during ageing.65 Duringprecipitation hardening, MgZn’ (or metastable �1’ phase)forms as rods parallel to the c-axis of the HCP unit cellwhile metastable �2’ phase forms as discs parallel to the(0 0 0 2) basal plane of the HCP unit cell.76–80 It hasbeen reported that the �1’ phase has a monoclinic struc-ture similar to that of Mg4Zn7 in aged Mg-8 wt.% Znalloy.76 Also, the yield strength of quasicrystalline par-ticle reinforced Mg–Zn–Y and Mg–Zn–Y–Zr magnesiumalloys was observed to increase with the volume fractionof the quasicrystalline phase based on the strengtheningeffect of the quasicrystalline particles.81 The icosahedralparticles in the Mg–Zn–Y alloy have been observed tobe stable against coarsening during elevated temperaturedeformation due to low particle-matrix interfacial energy.81

Concerning the aluminium-zinc or AZ series of magnesiumalloys, these alloys are characterized by:(a) low cost,(b) ease of handling,(c) good strength and ductility and(d) resistance to atmospheric corrosion.65 Such character-istics enable the common use of AZ series magnesiumalloys in weight-critical applications.65

Mg-RE (Rare Earth) alloys are the latest metallurgicalcategory of Mg alloys being developed. Currently, Mg-Yand Mg-RE alloys are in development for even more spe-cific metallurgical advantages but at generally higher costcompared to Mg–Zn and Mg–Al alloys.65 Compared to therest of the conventional melt-processed magnesium alloyfamily (mainly Mg–Al and Mg–Zn based alloys), it iswell-known that Mg–RE alloys have better elevated tem-perature strength retention characteristics.65 This is duemainly to the higher melting point of the Mg-RE inter-metallics present, meaning that these intermetallics com-mence going into solution only at higher temperatures.However, it is also well-known at present that long periodstacking/ordered phases (LPSO phases) also reinforce (inatom percent) Mg97Y2Zn1 type Mg-RE alloys.82 Here, theLPSO phases are generally composed of a solid solutionof Y and Zn atoms placed orderly in long periods alongthe Mg basal plane.83 Compared to the temperature whereMg-RE intermetallics start going into solution, the LPSOphase starts losing order at higher temperatures closer tothe liquidus temperature of the alloy.84 This implies that

the LPSO phase has even higher thermal stability com-pared to the Mg-RE intermetallic. The LPSO phase isalso more chemically homogenous (down to atomic lev-els) compared to the Mg-RE intermetallic region of similararea, necessarily meaning that the LPSO phase is morecorrosion resistant than the Mg-RE intermetallic region.Elevated temperature loss of strength and room tempera-ture corrosion susceptibility are major durability and costrelated disadvantages concerning magnesium alloy compo-nent usage that LPSO phase formation in selected Mg-REalloys may negate. Mg–Gd-Y alloys are also known forgood response to age hardening.85 Mg–Gd-Y-Zr alloy hasalso been co-rolled with AA7075 series aluminium alloy(as cladding) towards good tensile response.86

Table VII lists the mechanical properties of wroughtand cast aluminium-zinc (AZ), zinc-zirconium (ZK), rareearth (RE) and other magnesium alloys. Compared withTables I–V, based on basic mechanical properties, mostMg alloys listed in Table VII may be used to substitutecortical bone, may not be used to substitute cancellousbone, may be used to substitute enamel and dentin, maybe used to selectively substitute lower limb ligaments andtendons, and may be used in limited manner to substituteselected upper limb and trunk ligaments and associatedtissues.Concerning corrosion resistance, it is well known that

dissolved Cu readily binds with precipitating intermetal-lic phases in Mg alloys.65 Such binding tends to regu-late the precipitating intermetallic phases towards finersize, and increase the thermal stability of the intermetal-lic phase against dissolution into the alloy matrix, aswell as corrosion resistance. A good example is ZC71(Mg–6.5Zn–1.25Cu) where Cu mainly binds with Mg–Znintermetallics.65 Figure 1 shows the biodegradation char-acteristics of Mg alloys potentially suitable for clinicalapplications.96�100 Under physiological conditions (withCl− present), pure Mg rapidly corrodes with the followingmechanism:

Anodic reaction � Mg→Mg2++2e−

Cathodic reaction � 2H2O+2e− →H2 ↑ +2OH−

Mg2++2OH− →Mg�OH�2 ↓If phosphate and calcium ions are available, than Mg2+

may aid in the precipitation of calcium phosphate:

PO3−4 +Ca2++Mg2+ →MgxCay�PO4�

Mg(OH)2 forms a stable protective layer on the surface ofmagnesium implants in high pH (>11.5) environments, butlower pH (<11.5) will facilitate corrosion of magnesiumalloys in aqueous solution (see Fig. 2).101 The local pH atthe implant-bone interface is about 7.4 due to secondaryacidosis resulting from metabolic and resorptive processesafter surgery.72 The Mg(OH)2 layer is unable to adhere

Rev. Adv. Sci. Eng., 4, 1–18, 2015 7


Table VII. Mechanical properties of wrought and cast aluminium-zinc (AZ), zinc-zirconium (ZK), rare earth (RE) and other magnesiumalloys.65� 73� 87–99

Tensile yield Ultimate tensile Compressive yield Ultimate compressiveMaterial strength (MPa) strength (MPa) strength (MPa) strength (MPa)

Wrought AZ alloysRolled AZ31 [73] 150–220 255–290 110–180 –Extruded AZ31B/C-F [65] 200 255 97 –AZ31-GAE [73] 424 445 – –Extruded AZ61A-F [65] 205 305 130 –Extruded AZ91Da� b [87, 88] 215 296 – –Extruded AZ91Da� c [88] 226 313 – –AZ91+2Ca-GAE [73] 427 452 – –

Wrought ZK alloysExtruded ZK21A-F [65] 195 260 – –Extruded ZK31-T5 [65] 210 295 – –Extruded ZK40A-T5 [65] 255 275 – –Extruded ZK60Ad [89] 182±4 271±1 93±8 498±16Extruded ZK80Ad [90] 311±5 360±5 162±16 541±1

Wrought RE alloysExtruded AZR6-1e [91] 242 328 – –Extruded AZR6-2e� f [91] 286 352 – –LAE442 [73] 148 247 – –LANd442 [92] NA NA NA NALACe442 [93] NA NA NA NAWE43A-T6 [73] 162 250 – –WE43B [73] – 220 – 345Extruded WE43 [73] 198 277 – –Extruded MgYREZr g [94] 260 290 – –WZ21 [95] 150 250 – –

Other wrought Mg alloysMg0�8Ca [73] – 428 – –Mg(0–4)Ca [73] – 210–240 – –AM50A-F [73] – 210 113 –AM60B-F [73] – 225 130 –Extruded Mg–Zn–Mn [96] 246.5±4.5 280.3±0.9 – –ZEK100 [97] NA NA NA NAZX50 [95] 210 295 – –AX30 [98] NA NA NA NA

Cast AZ alloysSand cast AZ63A-T6 [65] 130 275 130 –Sand cast AZ81A-T4 [65] 83 275 83 –Sand cast AZ91C/E-T6 [65] 145 275 – –AZ91+2Ca [73] 147 – – –Sand cast AZ92A-T6 [65] 150 275 – –

Cast ZK alloysSand cast ZK61A-T5 [65] 185 310 – –Sand cast ZK61A-T6 [65] 195 310 – –Cast RE alloysSqueeze cast RZ5h [99] – – – 308

Other cast Mg alloysMg [91] 20.9±2.3 86.8±2.5 – –

Notes: aHot extruded at 250 �C. bRheocast material prior to extrusion. cDie-cast material prior to extrusion. dHeat treated at 150 �C for 1 hour. eAZR6 has nominalcomposition Mg94Al2Gd2Y1Zn1 in at%. fHeat treated at 200 �C for 24 hour. gMgYREZr has nominal composition 1.5–5 wt%Y, 2.5–5 wt%RE (rare earth metals), 0.1–2.5 wt%Zr, 0.01–0.8 wt%Zn, <0.5 wt%Al, <1.0 wt%Impurities, balance Mg. hRZ5 has nominal composition of 4.2 wt%Zn, 0.35 wt%Zr, 1.3 wt%RE (rare earth metals),balance Mg.

to and protect the implant surface under such conditions.Constant exposure to high chloride containing electrolyteof the physiological system causes accelerated invivo cor-rosion of the Mg implant. From Figure 1, it is apparentthat selected Mg alloys containing zinc (Zn), calcium (Ca)

and/or RE elements experience relatively lower corrosionrates under physiological conditions invitro. Comparingthe alloying elements on a binary system basis, gadolinium(Gd) is most effective in lowering the corrosion rate ofMg. The effectiveness of the alloying element in reducing

8 Rev. Adv. Sci. Eng., 4, 1–18, 2015


-0.5

0

0.5

1

1.5

2

2.5

Cor

rosi

on r

ate

(mm

/y)

OR

hyd

roge

nev

olut

ion

(ml/c

m2 /

day)

Mg-3.09Nd-0.22Zn-0.44Zr Mg-1Zn-1.1Mn-1Ca as-cast Mg-6Zn as-extruded

WE43 as-cast Mg-Si as-rolled Pure Mg as-rolled

Mg-1Zn as-rolled Mg-5.6Zn-0.55Zr-0.9Y Mg-1Ca as-extruded

ZK60A Mg-2Sr as-rolled Mg-0.8Ca as-extruded

Mg-Zn-Mn as-cast WE43 as-extruded

in SBF in Hank’ssolution

invivo in SBF-H2 in Hank’ssolution-H2

Mg-Ca Mg-Sr Mg-Zn Mg-Gd Mg-Dy Mg-Y Mg-Sn Mg-Si

0

5

10

15

20

25

0 5 10 15 20 25

Alloying elements (wt%)

Deg

rada

tion

rate

(m

m/y

)

Fig. 1. Biodegradation characteristics of Mg alloys potentially suitable for clinical applications.96� 103 Shaded regions indicate desired corrosion ordegradation rates.

the corrodibility of Mg reflects the stability of the alloyingelement oxide in aqueous media.

The stability of the alloying element oxide in aque-ous media reinforces the logic behind which Mg alloy touse for clinical osteo-applications, depending on intendedduration of patient recovery. Apparently, selected Mgalloys containing zinc (Zn), calcium (Ca) and/or RE ele-ments are best suited for longer durations of patient recov-ery while other Mg alloys in the same family are bestsuited for relatively shorter durations of patient recovery.However, there is no clear opinion at present regarding thetoxicity levels of many alloying elements in Mg alloys.

Regardless of the duration taken by the Mg alloyimplant to corrode invivo, it is desired that the corrosion

occurs in a stable and predictable manner, like in uniformcorrosion. This is best achieved by including an alloyingelement whose native oxide film easily forms and adheresto the magnesium substrate without significantly deterio-rating and flaking off. This is also best achieved by includ-ing an alloying element which forms fine second phaseprecipitates with the Mg matrix. Second phase precipitatesare cathodic respective to the alloy matrix and coarser pre-cipitates contribute to increased micro-galvanic corrosion.This is where RE elements have a significant advantageover other alloying elements. RE elements are known toform stable oxide films easily and also form fine secondphase precipitates. Syntellix AG and the Hannover Medi-cal School in Germany have designed the MgYREZr based

Rev. Adv. Sci. Eng., 4, 1–18, 2015 9


E(V)

Mg2+

Mg

Mg(OH)2

pH

3.5

0.25

–1.00

0.10

–2.40–2.55

Fig. 2. Pourbaix (potential-pH) equilibrium diagram for magnesium-water system at 25 �C.101

Magnezix cannulated compression screw for orthopaedicapplications where a Phase I clinical trial has been recentlysuccessful.94�102 It is unclear if any surface coating (e.g.,MgF2, apatite, carbonate etc.) exists for the Magnezixmedical device. The nature of RE element used alsoneeds more discussion. LAE442 is another Mg-RE alloywhere more stable corrosion has been observed invivo.73

However, a mix of rare earths known as mischmetal ordidymium was used. Mischmetal and didymium can eachdiffer in composition depending on the time and/or date ofpurchase. Mischmetal is cerium (Ce) based with additionsof lanthanum (La) and neodymium (Nd) while didymiumis Nd based with addition of praseodymium (Pr).65�103 Forreliable reproduction of this Mg-RE alloy, Ce or Nd wereused to primarily replace the mischmetal or didymiummixtures.92�93 Interestingly, LANd442 exhibited stable cor-rosion characteristics (however not as stable comparedto LAE442) while LACe442 exhibited unstable corrosioncharacteristics.Regarding toxicity levels of RE elements, short-term

effect of RE elements invitro were evaluated.104 Single REelements were added to different cell lines and the effectswere assessed. Differences between light (La, Ce, Pr) andmedium to heavy (Nd, Europium (Eu), Gd and Dyspro-sium (Dy)) RE elements were observed. Generally, lightRE elements showed toxic effects at lower concentrationsand it was concluded that La and Ce should be used onlywhen absolutely necessary. Concerning long term effectsof RE elements invitro as well as invivo, no systematicstudies or trials have been completed. Significant RE expo-sure and effects among miners in RE mines located withinMongolia (China) have been documented.105 In the humanbody, RE elements are known to accumulate in the liver,kidneys and bone without being metabolised.

5. BIODEGRADABLE CERAMICSBone is essentially composed of calcium based phosphatesas the mineral phase and collagen as the organic bindingphase. Calcium phosphate based bioceramics are the maingroup used clinically as biodegradable ceramics. Bioce-ramics are used to fill tooth and bone defects, fix bonegrafts, fractures and prostheses, and replace diseased (e.g.,cancerous) tissue. Based on reaction with host tissue, bio-ceramics are bioactive and most are resorbable, encourag-ing and allowing replacement by bone and organic tissueat the site of injury. The similarities between the bone min-eral phase and the structural and surface features of thebioceramics enable binding to bone without formation ofundesirable fibrous connective tissue interface.106�107

The most commonly used bioceramics are hydrox-yapatite (HA, Ca10(PO4�6(OH)2�, �-tricalcium phosphate(�-TCP, Ca3(PO4�2), derivatives as well as combinations.These are crystalline bioceramics. Densified HA is chemi-cally most stable and the least resorbable invivo comparedto the other bioceramics. While the other bioceramics arecompletely resorbed after implantation into the injuredbone site, densified HA may remain integrated into theregenerated bone or slightly resorbed over time.108�109

There is also a range of calcium phosphates that areinjectable and harden within the bone cavity withoutreleasing too much heat and causing premature death ofsurrounding cells.110�111 HA has been used to fill smallbone defects after bone tumour resection or after bone lossin fresh fractures but there has been comparatively lessindication of HA being used to fill large bone defects.112

Slow degradation of HA and minimal resorption after a12 week implantation period in rabbit’s femora has alsobeen reported.113

Compared to crystalline bioceramics such as HA,�-TCP, derivatives as well as combinations, there arealso amorphous bioglasses and glass-ceramics comprisedmainly of CaO, Na2O, SiO2 and P2O5.

114–116 The mainadvantage bioglasses and glass-ceramics have over crys-talline bioceramics is that even though they are mechani-cally weaker and resorb faster than crystalline bioceramics,they form a dense and strong interface with natural boneup to 3 times faster than crystalline bioceramics.117 Sili-con (Si) in biological ceramics and glasses has a significanteffect in the bone regeneration process. Compared to non-doped apatites, Si has been incorporated into apatites toinduce the formation of a higher amount of bone tissue.118

Here, Si was observed to improve bioactivity of apatite byforming Si–OH groups on the material surface. The Si–OHgroups triggered the nucleation and formation of apatitelayers on the surface improving the material-bone bond-ing characteristics. Generally, SiO2 is the network formerwhile CaO, Na2O and P2O5 act as modifying agents inbioglasses and glass-ceramics. Table VIII lists the physi-cal characteristics, mechanical properties and compositionof biodegradable ceramics and glasses,21 while Figure 3

10 Rev. Adv. Sci. Eng., 4, 1–18, 2015


Table VIII. Physical characteristics, mechanical properties and composition of biodegradable ceramics and glasses.21

Ceramics Bioglass Glass-ceramic

Property HA TCP 45S5 S45PZ Ceravital Cerabone A/W I1mappant Biovert

Density (g/cc) 3.16 3.07 2.66 – – 3.07 – 2.8Hardness (HV) 600 – 460 – 680 – 500 460Compressive strength (MPa) 500–1000 460–680 – – 500 1080 – 500Bending Strength (MPa) 115–200 140–154 110–140 – – 215 160 500Young’s modulus (GPa) 80–110 33–90 35 – 100–150 218 – 70–88Index of Bioactivity IB 3.1 – 12.5 3.8 5.6 7.5 – –

Composition of biodegradable ceramics and glasses

Oxide/phase HA TCP 45S5 S45PZ Ceravital Cerabone A/W I1mappant Biovert

Na2O – – 24.5 24 5–10 0 4.6 3–8K2O – – 0 – 0.5–3.0 0 0.2 3–8MgO – – 0 – 2.5–5.0 4.6 2.8 2–21CaO – – 24.5 22 30–35 44.7 31.9 10-34Al2O3 – – 0 – 0 0 0 8–15SiO2 – – 45.0 45 40–50 34.0 44.3 19–54P2O5 – – 6.0 7 10–50 16.2 11.2 2–10CaF2 – – 0 – – 0.5 5.3 3–23B2O3 – – 0 2 – – – –Phase Apatite Witlokite Glass Glass Apatite glass Apatite �- Apatite �- Apatite phlogopite

wollas-tonite wollas-tonite glass

shows the compositional dependence of bone bonding toglass.21 The index of bioactivity or IB listed in Table VIIIis a measure of how active the bioceramic is toward denseinterphase formation with natural bone at the site of injury.Briefly, IB indicates how fast (relatively) the bioceramicmay be capable of forming the dense interphase with natu-ral bone. Due to lower crystalline form, the bioglasses andglass-ceramics have higher IB than crystalline bioceramics,but tend to be mechanically weaker than the crystallinebioceramics as listed in Table VIII.

Compared with Tables I–V, based on basic mechan-ical properties, densified HA, TCP, Ceravital, CeraboneA/W and Biovert (as listed in Table VIII) will presentimmediate stress shielding effects when implanted into

A-Bone bonding

B-Bioinert

C-Total dissolution

D-Non glass forming

S-Osteoinduction(shaded)

6% P2O5

CaO

SiO2

Na2O

A

B

C

D

S

IB = 0

IB = 2IB = 5 IB = 8 IB = 10

Fig. 3. Compositional dependence of bone bonding to glass. Index ofbioactivity (IB) contours are superimposed onto the phase diagram.21

cortical bone, cancellous bone, enamel and dentin, lowerlimb ligaments, tendons and fascia, and upper limb andtrunk ligaments and associated tissues. Comparatively,45S5 bioglass should not present immediate stress shield-ing effects. Most bioceramics (in densified form) listed inTable VIII may be used to substitute cortical bone, maynot be used to substitute cancellous bone, may be used tosubstitute enamel and dentin, may be used to selectivelysubstitute lower limb ligaments and tendons, and may beused in limited manner to substitute selected upper limband trunk ligaments and associated tissues. The IB valueof the bioceramic reinforces the logic behind which bio-ceramic to use for clinical osteo-applications, dependingon intended duration of patient recovery. Crystalline bio-ceramics like HA and TCP having relatively lower IB arebest suited for longer durations of patient recovery whileamorphous glasses like 45S5 and S45PZ and glass ceram-ics like Ceravital, Cerabone A/W, I1mappant and Bioverthaving relatively higher IB are best suited for relativelyshorter durations of patient recovery. This is especially thecase even when the non-densified forms of bioceramics,such as porous structures or powders having higher spe-cific area per unit volume for chemical interactions, areinserted into bone cavities for repair.MgO, TiO2 (anatase) and K2O have also been used

as network modifiers in amorphous bioglass and glassceramic.119–121 Bioglass surfaces have also been mod-ified to further enhance their bioactivity by coat-ing them with adhesive proteins to promote celladhesion.122 Clinical uses of amorphous glass and glassceramic so far include filling materials in benign tumorsurgeries,123 reconstruction of defects in facial bones,124�125

Rev. Adv. Sci. Eng., 4, 1–18, 2015 11


tympanoplastic reconstruction,126 lumbar fusion,127 treat-ment of periodontal bone defects,128�129 repairing orbitalfloor fractures,125�130 obliteration of frontal sinuses,131–133

reconstruction of the iliac crest defect after bone graftharvesting,134 and repairing orbital floor fractures.125�130

Reports on clinical applications are found for Depuy SpineConduit® (TCP), Medtronic MasterGraft® (HA, TCP),Stryker Vitoss® (TCP), Synthes ChronOS®, Norian SRS®

(calcium phosphate), Stryker Cortoss® (bioactive glass),etc. Synthetic HA has good compatibility with tissue but isclinically limited due to moderate resorbability invivo andmechanical properties that differ from surrounding tissuesand bones.135 With the aim of enhancing mechanical prop-erties and bioactivity, there have been efforts to dope HAwith magnesium (Mg),115 strontium (Sr),136 silicon (Si),137

and carbonate (CO2−3 ).138 HA has also been doped with

silver (Ag) for curing infected bone defects.139

6. BIODEGRADABLE COMPOSITESBioceramics and biodegradable polymers have beencombined into biodegradable composite pastes thatmay be injected into bone cavities for orthopedicapplications.140–156 Hydroxyapatite (HA) and tricalciumphosphate (TCP) have improved the osteoconductivityand bone bonding of the host polymer.140–143 Since HAand TCP resorp in basic conditions, any acidic degra-dation products released during polymer matrix degrada-tion may be pH buffered, enabling foreign body reactionminimization.144�145 The ceramic reinforcement may alsoact as a hydrolysis barrier, further delaying degradationof the polymer matrix.146 Interestingly, auto-generatedincrease of local acidity due to degradation of poly(lacticacid) (PLA) for example may enhance solubility of theHA or TCP ceramic that may be used in new boneformation.147 Polycaprolactone (PCL) has also been com-bined with calcium phosphate based reinforcement towardthe same effect.148 Concerning mechanical properties,there have been reports of calcium phosphate based rein-forcements increasing stiffness and strength of the hostpolymer while also altering the rheological or viscousflow properties of the host polymer intended for injectioninto bone cavities.149–152 Fast setting calcium phosphatecements (CPCs) are also another example of biodegrad-able composite.153�154 These cements set within the bonecavity when injected near body temperature and avoid ther-mally damaging surrounding tissue. Chitosan has signifi-cant pharmacological benefits for bone formation and hasbeen reported as a component of CPCs.155 Poly(propylenefumarate) (PPF) was developed with TCP and CaCO3

fillers. This injectable composite hardened in 24–36 hrs,had mechanical properties much higher than human can-cellous bone, and was sufficiently workable to be packedinto bone defects of complex shape.156

Apart from injectable polymer-based or ceramic-basedbiodegradable composites, biodegradable composites have

been used in medical applications where stiffness andstrength is low, e.g., tissue scaffolds. However, concerningsignificant clinical application for implantable biodegrad-able material systems, the ability to bear appropriate loadat least in the initial stages of recovery prior to signifi-cant resorption and load transfer to new tissue is critical.This is exactly where biodegradable Mg alloys or com-posites have significant potential for clinical application.Table IX lists the mechanical properties of wrought, sin-tered and cast aluminium-zinc (AZ), zinc-zirconium (ZK),rare earth (RE) and other magnesium nanocomposites andmicrocomposites.89�91�157–178 Compared with Tables I–V,based on basic mechanical properties, most Mg alloynanocomposites and microcomposites listed in Table IXmay be used to substitute cortical bone, may not be usedto substitute cancellous bone, may be used to substi-tute enamel and dentin, may be used to selectively sub-stitute lower limb ligaments and tendons, and may beused in limited manner to substitute selected upper limband trunk ligaments and associated tissues. An attrac-tive feature of wrought Mg alloy nanocomposites is thehigh ductility (>15%, >20%, >25%) in selected formu-lations. The high ductility may permit cold working tobe performed to further increase strength at the expenseof some of the excess ductility. However, it is probable(though not confirmed) that such further strengthening mayworsen the corrosion behavior invivo. In the as-extrudedor stress-relaxed state, it has been often shown that thenanoparticle additions to the different Mg alloys regu-late the precipitation of second phase particles towardsnanosize.89�91�157–171 This breaks up surface cathodic sitesand overall improves corrosion resisstance.179 Concern-ing Mg microcomposites, it has been reported that themicroparticle reinforcement deteriorates the invitro corro-sion characteristics of the composite.172–178 Another issuewith the Mg alloy nanocomposites is the potential cyto-toxicity of the exsitu added and/or second phase nanopar-ticles that may enter the body’s circulatory system duringinvivo degradation of the biodegradable implant.180 Thesenanoparticles may accumulate over time in different partsof the body, and because of the abundant and activatedtight surface area inevitably interfere with selected cellbased functions, or even affect genetic related material inthe body causing hereditary changes. Provided the totalconcentration of nanoparticles in the alloy nanocompositeis 3 vol% and the nanoparticles are about 50 nm in diam-eter and well dispersed into the body’s circulatory systemduring implant degradation, assuming 2 years completedegradation time for a 10 cm long alloy nanocompositerod of 8 mm diameter that is inserted at the damaged tis-sue site, as many as 2�304× 1017 or 3�826× 10−7 molesof nanoparticles may be in circulation throughout thecourse of the 2 year recovery time. Like the rare earthalloying elements discussed earlier in the biodegradablemetal (magnesium) section, the nanoparticles may remain

12 Rev. Adv. Sci. Eng., 4, 1–18, 2015


Table IX. Mechanical properties of wrought, sintered and cast aluminium-zinc (AZ), zinc-zirconium (ZK), rare earth (RE) and other magnesiumnanocomposites and microcomposites.89� 91� 157–178

Material Tensile yield strength (MPa) Ultimate tensile strength (MPa) Tensile ductility (%)

Wrought AZ nanocompositesExtruded AZ31/1.5vol%Al2O3 [157] 204±8 317±5 22.2±2.4Extruded AZ31/1.0vol%CNT [158] 190±13 307±10 17.5±2.6Extruded AZ31/AA5083/1.0vol%CNT [159] 221±4 321±1 12.0±1.0Extruded AZ31/AZ91/1.5vol%Al2O3 [160] 232±13 339±10 15.9±0.5Extruded AZ31/AZ91/1.5vol%TiC [161] 236±8 337±7 14.5±0.7Extruded AZ31/AZ91/1.5vol%Si3N4 [162] 232±2 331±2 13.1±0.5Extruded AZ81/1.5vol%Al2O3 [163] 212±12 339±11 13.1±1.0Extruded AZ81/1.5vol%CNT [164] 209±9 328±4 13.7±2.2Extruded AZ81/1.5vol%Si3N4 [165] 229±9 328±7 10.7±0.9Extruded AZ91/ZK60A/1.5vol%AlN [166] 236±6 336±4 13.8±1.0Extruded AZ91/ZK60A/1.5vol%B4C [167] 222 345 25.1Extruded AZ91/ZK60A/1.5vol%B4C CW a [167] 370 510 12.2

Wrought ZK nanocompositesExtruded ZK60A/1.5vol%Al2O3 [89] 175±2 305±2 18.1±0.9Extruded ZK60A/1.0vol%CNT [168] 180±6 295±8 15.0±0.7Extruded ZK60A/1.5vol%Si3N4 [169] 198±6 313±4 12.2±0.8Extruded ZK60A/1.5vol%TiC [170] 184±2 309±3 11.6±1.4Extruded ZK60A/1.5vol%SiB6 [171] 175 309 22.7Extruded ZK60A/1.5vol%SiB6 CWa [171] 340 478 11.3

Wrought RE nanocompositesExtruded AZR6/1.5vol%ZrB2-1

b [91] 270 362 12.2Extruded AZR6/1.5vol% ZrB2-2

b� c [91] 307 375 7.3Extruded AZR6/1.5vol% ZrB2-3

b� c� d [91] 407 463 4.3

Sintered AZ microcompositesAZ91D/20HA [172] 264.3 – –AZ91/30FA [173] 112.4 – 4.51

Sintered ZK microcompositesZK60/10CPP [174] 215 230 –ZK60/20CPP [174] 210 220 –ZK60/30CPP [174] 195 200 –

Other Sintered Mg microcompositesMg/10HA [175] 117.3 171.6 6.7Mg/20HA [175] 105.8 146.9 4.3Mg/30HA [175] 71.7 92.1 2.6Mg/40HA [175] – 68 –Mg/1Ca [176] 147.78 217.28 14.36Mg/5Ca [176] 183.32 202.72 9.03Mg/10Ca [176] 119.63 200.25 7.74

Other cast Mg microcompositesMg/5HA [177] 122.3 171.1 –Mg/10HA [177] 137.0 146.4 –Mg/15HA [177] 129.6 136.7 0.3ZM61/15HA [178] 225.5 225.5 0.3

Notes: CNT: Carbon NanoTube, HA: HydroxyApatite, FA: FlouroApatite, CPP: Calcium PolyPhosphate. aHypothetically 10% cold-worked. bAZR6 has nominal compositionMg94Al2Gd2Y1Zn1 in at%. cHeat treated at 200 �C for 24 hour. dWith nano-LPSO-grain clusters as reinforcement.

unmetabolized in the body in the long term, enabling var-ious acute or chronic effects.

7. FUTURE DIRECTIONSFuture directions of this field involving biodegradablematerials include developments on cellulose. Bacterial cel-lulose has been physically modified to guide calciumphosphate precipitation from simulated body fluid.181�182

Highly porous composite scaffolds for trabecular bone

regeneration have also been successfully synthesized usinghydroxypropyl-methyl-cellulose.183 Bamboo pulp (as areliable source) has been mechanically defibrillated toyield cellulose nanofibrils for medical applications.184

Cellulose acetate nanofibers have also been electro-spun into functional webs, where pH and hydroxyapatitenanoparticle content were each varied to enhance invitrobioactivity in simulated body fluid.185 Cellulose containslignin which may prove challengable to electrospin unless

Rev. Adv. Sci. Eng., 4, 1–18, 2015 13


mixed in a solution with polyacrylonitrile.186 Well dis-persed lignin nanoparticles have also been synthesized inan environmentally friendly and upscalable manner usingacid precipitation technology.187

Directly relevant to bone tissue regeneration and engi-neering, the evolution of scaffolds as bone graft substitutesin the process of recreating the bone tissue micro-environment, including biochemical and biophysical cueshas been reviewed.187 The most common form of bonegraft is the autograft but its use can lead to complica-tions such as pain, infection, scarring, blood loss, anddonor-site morbidity. The alternative is allografts whichlack the osteoactive capacity of autografts and carry therisk of carrying infectious agents or immune rejection.Other approaches, such as the bone graft substitutes,have focused on improving the efficacy of bone grafts orother scaffolds by incorporating bone progenitor cells andgrowth factors to stimulate cells.187 Also, the microstruc-tures and phase compositions of artificial and bone-derivedhydroxyapatites (HAs) have been compared.188 Bone-derived HA was observed to consist mainly of HA and asmall amount of MgO. Hot-pressed HA compacts showedhomogeneous microstructures and densities of 95–97%but grain sizes and microstructures varied with the start-ing powders. Self-assembling peptide nanofiber scaffoldshave also been developed for bone tissue engineering.189

Concerning resorbability of HA, barium and fluorine havebeen observed to increase the invitro resorbability ofsynthetic HA.190 The magnetic and morphological prop-erties of ferrofluid-impregnated hydroxyapatite/collagenscaffolds have also been reported, where sufficient mag-netization was used to attract potential magnetic carri-ers for transporting bioactive agents which favoured boneregeneration.191 The osteoblast attachment on HA scaffoldshas also been increased by performing an air atmosphericpressure plasma jet pretreatment for drop-wise loading ofdexamethasone.192

Concerning magnesium alloys, addition of calcium (Ca)to increase tensile yield strength to 340 MPa based onincreased hot rolling speeds (and recrystalization) has beenreported.193 For corrosion resistance, improved wear per-formance and reduced thermal transport behavior on thesurface, the formation of nanostructured oxidic coatingson Mg alloys using plasma electrolytic oxidation (PEO)has also been reported.194 Also, polymer encapsulation viaRF plasma polymerization has been critically discussedtowards tailoring biodegradability and biocompatibility ofmagnesium.195

Nanoparticles and nanotechnology continue to fuel thefuture of many fields including this one. Polyurethanenanofibers containing aqueous extract of Grewiamol-lis Juss have been produced using electrospinning forenhanced antimicrobial effects against Escherichia coliATCC 52922 (Gram negative) and Staphylococcus aureusATCC 29231 (Gram positive) in liquid medium.196 Con-ductive polymer nanofibers have also been electrospun into

scaffolds for the pursuit of biomimetic extracellular matrix(ECM) structures for adhesion, proliferation, and differ-entiation of cells involving electrical stimuation.197 Tita-nia and polylactide (including composite) nanoparticles aswell as fucoidan-chitosan nanospheres have been reportedto have significant potential for drug delivery applicationsbased on good biocompatibility and biodegradability.198�199

Pyrolysis derived carbon matrix nanocomposites havealso been observed to be sufficiently protective of mag-netic nanoparticles (embedded) for biomedical applicationinvivo.200 Carbon nanotubes (CNTs) have been observed toincrease cellular uptake of three-dimensional biomimeticcollagen hydrogels.201 Polycaprolactone scaffolds havealso been critically reviewed.202 Here, the incorporationof nanofillers or blending of polycaprolactone with otherpolymers has yielded a class of hybrid materials with sig-nificantly improved physical and chemical properties suchas strength, porosity, microstructure, controllable degra-dation rates, and bioactivity that are important for tissueengineering. Concerning potential neuroprotection usingCerybrolysin Therapy, the role of functionalized magneticiron oxide nanoparticles in the central nervous system(CNS) injury and repair has been studied.203 The mag-netic iron oxide nanoparticles were observed to be safe forthe CNS in disease conditions when co-administered withcerebrolysin. The interaction of nanomaterials with humancells is the most important criterion for biomedical appli-cation and has been studied in detail.204 Three commonapproaches have been suggested for nanomaterial mecha-nisms for cellular interaction and internalization: direct dif-fusion or disruption to the plasma membrane, endocytosis,and entry through ion channels and transporter proteins.The antimicrobial properties of nanobiomaterials contain-ing silver and its salts have also been reported.205�206

As a form of soft tissue engineering, vascular tissueengineering is tied in with hard tissue engineering formssuch as bone tissue engineering concerning future trends.Polycaprolactone (PCL) micro/nanofibers have been elec-trospun into 2D scaffolds where cell culture results showedthat the cells are not only attached to the scaffold butalso integrated with it i.e., the cells are embedded into thePCL scaffold.207�208 Small deformable Janus particles withzwitterionic nature smoothly transport through cathetersand slowly form large aggregates with effective vascu-lar occluding properties. In arterial embolization, embolicparticles should be deformable for smoother injection andlarge enough for effective occlusion of arteries. The uniqueassembly behavior of Janus particles have been reportedto possess significant potential in exceeding the contem-porary limits of embolization technology.209

8. CONCLUSIONS1. On the basis of mechanical property comparison, eachclass of densified material (polymer, metal (magnesium),ceramic, composite) may be used to substitute cortical

14 Rev. Adv. Sci. Eng., 4, 1–18, 2015


bone, may not be used to substitute cancellous bone, maybe used to substitute enamel and dentin, may be used tosubstitute lower limb ligaments and tendons, and may beused to substitute selected upper limb and trunk ligamentsand associated tissues.2. Concerning longer times of recovery from more severeorthopedic trauma, PLLA (poly L-lactic acid) and TMC(trimethyl carbonate) polymers, selected Mg alloys con-taining zinc (Zn), calcium (Ca) and/or rare-earth (RE)elements, or hydroxyapatite (HA) and tricalcium phos-phate (TCP) as crystalline bioceramics apparently are mostsuitable.3. Concerning biodegradable alloys, Mg alloys or com-posites have significant potential for clinical applicationwhere the ability to bear appropriate load at least in theinitial stages of recovery prior to significant resorption andload transfer to new tissue is critical. However, there isno clear opinion at present regarding the toxicity levelsof many alloying elements in Mg alloys, particularly con-cerning RE elements. There is also the potential issue ofnanoparticle cytotoxicity concerning alloy nanocompositedegradation invivo.4. Concerning bioceramics, the higher index of bioactiv-ity (IB) of amorphous glasses and glass ceramics reflectsthe superior ability to form a dense interphase and bondstrongly with bone, despite mechanical inferiority to crys-talline bioceramics.5. Concerning biodegradable composites, the combinationof bioceramic and biodegradable polymer is synergisticbased on direct improvement of mechanical property anddegradation resistance of the polymer, indirect reduction offoreign body interactions, and direct increase in toughnessof the ceramic.6. Future directions in this field include developmentson cellulose as a biodegradable material, bone tissueregeneration and engineering, stronger and more corro-sion resistant and biocompatible magnesium alloy sys-tems, applicable nanoparticles and nanotechnology, andsoft tissue engineering such as vascular tissue engineering.

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