Atomic force microscopy and indentation force measurement of bone

Advanced Review

Atomic force microscopy andindentation force measurementof bonePhilipp J. Thurner

This review is summarizing the results obtained from atomic force microscopy(AFM) and nanoindentation experiments to date. The combination of bothtechniques is especially powerful. It allows to carefully choose indentation locationsas well as the post-hoc analysis of the created indents, and hence the possibility toassess the properties of microstructural elements of bonessue. In addition, AFMhas improved our understanding of bone ultrastructure and force spectroscopyexperiments have led to the discovery of a molecular self-healing effect of bonethat may be based on a small fraction of unstructured proteins. Nanoindentationmeasurements on bone, pose inherent problems since bone is an anisotropicsolid showing elastic, viscoelastic, and time-dependent plastic behavior. Hence,derived parameters such as elastic modulus and hardness are to some extentdependent on measurement protocols. However, the development of extensions tothe OliverPharr method, being the most widely used analysis method, as well asnovel dynamic testing techniques could improve the situation. Nanoindentationis widely used to study bone tissue and some important principal findings havebeen reported to date. These are presented here together with specific resultsfrom nanoindentation experiments of human and animal bones and tables arepresented collating the data that can be found in the literature to date. 2009 JohnWiley & Sons, Inc. WIREs Nanomed Nanobiotechnol 2009 1 624649

Bone is a remarkable biological material providingskeletal stability and support and protection ofvital organs for a lifetime. Research on bone goes backfor more than a century and one of the most impor-tant aspects of bone, i.e., its capability to remodeland adapt to the loads it is subjected to had alreadybeen reported in the 19th century by Julius Wolff.1

Nevertheless, bone research is by far not at its end.Especially bone, as a material, due to its intricatehierarchical structure ranging from the molecular tothe macroscopic scale is still not very well under-stood. Mostly, this emerges from that fact that natureuses nanoscale components when building compos-ite materials compared to engineers. This approachseems to have important advantages. As pointed outbyGao et al.,2 suchmaterials are much less sensitive toflaws. Whereas this approach chosen by nature is truly

Correspondence to: [email protected] of Engineering Sciences, University of Southampton,Highfield, Southampton SO17 1BJ, UK

DOI: 10.1002/wnan.056

beneficial for healthy individuals, it also poses highdifficulties when determining structurefunction rela-tionships. However, with the ongoing developmentand improvement of apparatuses and experimentaltechniques such as the atomic force microscope (AFM)and nanoindentation, among others, the means ofprobing both nanoscale or ultrastructure and mechan-ical function of bone are steadily improving. Perhapsthe most important common aspect of these two tech-niques is that samples can be investigated in ambientenvironment or even in buffered saline liquids, veryclose to its physiological state. Ever since its inventionin 1986 by Binnig et al.,3 the AFM has been exten-sively used in biology and biophysics for imaging andmechanical probing of DNA, proteins, tissues, andbiomaterials. In principle the technique relies on asmall cantilever with a very sharp tip (typical radius 10 nm) that is scanned across a surface of inter-est with a controller using the detected deflection oramplitude, in the case of an oscillating cantilever,as a feedback signal for adjusting the height of thecantilever with respect to the sample surface. These

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WIREs Nanomedicine and Nanobiotechnology AFM and indentation force measurement of bone

height adjustments represents the sample topogra-phy. In addition to imaging, the AFM cantileverscan also be used as very sensitive force sensors, andhave been used in rather elegant ways to measure theforces involved in antibodyantigen interactions, theunfolding of proteins with globular domains, or themechanical behavior of polysaccharides.46 Nanoin-dentation emerged in the early 1980s from improvingmicroindentation methods by enhancing force anddepth sensing capabilities.7 It was originally devel-oped to investigate coatings, and it was over a decadebefore the first work on bone nanoindentation waspublished. Nowadays, nanoindentation is probablythe most important tool for the assessment of theelastic modulus of bone tissue. As mentioned before,bone is a nanocomposite and therefore of interestto the materials science community from a basic sci-ence point of view. More importantly, however, thereis the clinical need for better diagnostic tools andtreatments for the bone disease osteoporosis, whichis mostly affecting women after menopause and theaged, resulting in a dramatic increase in fracture risk.Fractures in turn result in decreased quality of life,pain, and even morbidity for patients8,9 and a largesocioeconomic burden for society.10,11 Osteoporosisstems from changes in cell signaling between osteo-clasts, osteoblasts, and osteocytes, which can resultin a multitude of changes within bone tissue. Wheninvestigating whole populations, the disease has aprominent effect on bone mass, which is officiallyused as the diagnostic parameter to determine fracturerisk.12 This is despite the better knowledge that forthe individual this diagnosis is rather ineffective,13,14

but on the other hand also the only tool availableto the clinician. Hence, to provide additional toolsa large number of investigations have been carriedout to characterize changes in the microarchitectureof trabecular bone, which also affect the mechanicalcompetence of bone.15 However, whereas bone qual-ity is usually mentioned as an important parameter,investigations into bone matrix material propertiesare more scarce and especially so in investigationwith the aim to determine structurefunction rela-tionships. Results from such studies are encouragingpointing to the fact that tissue composition doesindeed change,16 and that there can be dramaticchanges in bone properties.17 Hence, research andunderstanding of bone quality, and in this context,i.e., the structurefunction relationship between bonecomposition at the nanometer scale and mechanicalcompetence from the nanometer to the macroscopicscale, is still emerging and has not yet reached a levelthat can currently aid in diagnosing patients at risk

for osteoporotic fractures. The purpose of this com-munication is to review what has been achieved todate; the most important discoveries in bone ultra-structure and mechanical competence from AFM andnanoindentation studies will be discussed, and quan-titative results from both techniques reported to datewill be compared, data analysis and experimentalmethods will be discussed, and limitations and futuretrends and challenges will be pointed out.

BONE ULTRASTRUCTURE ASSESSEDWITH AFM

Bone is a hierarchically structured tissue withstructural features ranging from the nano- to themacroscale, which are all influencing the mechanicalbehavior of a whole bone or a macroscopic piece ofbone.

The Individual ComponentsAt the lowest hierarchical level, bone consists offibrils made of collagen type I, mineral crystals ofcarbonated hydroxyapatite, a variety of unstructuredso-called noncollagenous proteins (NCPs), and water.All these components are put together in an intricateway. Much of our understanding of the organizationof bone at this level originates from imaging studiesusing techniques such as scanning electronmicroscopy(SEM),18,19 transmission electron microscopy,20 elec-tron tomography21 as well as AFM. In addition,immunohistochemical22,23 as well as spectroscopictechniques24,25 have given us a great deal about bonecomposition. This review, however, is focused onstudies employing AFM and nanoindentation. Hence,I will first discuss AFM investigations on specificisolated components of bone and then show howfurther AFM studies have been useful to learn moreabout the assembly of these different components.

Collagen Type I FibrilsThe main structural protein in bone, similar to otherconnective tissues like tendon, ligaments, and skin,is collagen type I, which assembles into microfibrilsand in turn into fibrils with typical diameters from 50to 200 nm (Table 1). An intriguing feature of thesefibrils is that they show a distinct banding patternof about 67 nm in length (Table 1) when imagedwith AFM, usually called D-banding,26,27 as shownin Figure 1(a). As an explanation of this distinct pat-tern that has extensively been studied with AFM,among other methods, the model proposed by Hodgeand Petruska27,28 is cited to this date. These authors

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TABLE 1 Overview of Structural Parameters of Major Components of Bone UltrastructureCollagen Type I Fibrils and HydroxyapatiteCrystalsObtained from Atomic Force Microscopy Experiments

Collagen Type I Fibrils Hydroxyapatite Crystals

D-Banding Diameter Size Material Reference

6470 nm 50 nm to 1m Tissue engineered bone matrix 40Not observable 159 nm Demineralized cortical rat bone 41

67 nm Partially to fully demineralized humandentin

42

69 nm 225 96 nm Bovine cortical bone 4367 nm 0.821.4 m Elephant dentin (ivory)66.5 1.4 nm 62.2 7.0 nm Bovine cortical bone 4469.4 4.3 nm 141 12 nm Elephant dentin (ivory) 15 5 nm length, 10 3 nm

width, 0.61 0.19 nmthickness

Bovine femur and tibia (mature corticalbone, 48 year old animals)

45

67 nm 50200 nm 30120 nm diameter, 310 nmthickness

Bovine trabecular bone from vertebra 46

67 2 nm 90 20 nm Human trabecular bone from femoral head,69 year male osteoarthritic donor

47

6575 nm 95 15 nm 2060 nm diameter, 5 2 nmthickness

Ovine bone (Swiss alpine sheep) from adultfemale

48

67 2 nm,120 10 nm

110 10 nm 76143 nm diameter Bovine cortical bone from 20-month-oldanimal

49

9 3 nm length, 6 2 nm width,2.0 1.2 nm thickness

Bovine femur and tibia (young corticalbone, 35 week old animals)

38

67.9 nm 80105 nm Human dentin, 2234 years 50

67 nm 100 nm 47 18 nm length, 33 12 nmwidth

Bovine vertebral trabecular bone 51

66.79 1.25 nm 50250 nm diameter Bovine femoral trabecular and cortical bone 37

explained the observed periodicity by a quarter stag-gered, side-by-side alignment of five collagen triplehelices resulting in an overlap and a gap zone. How-ever, as more recently pointed out by Bozec et al.,

this 2D model has a number of shortcomings andmost importantly fails to explain how the D-bandingis preserved when going from five triple helices tocollagen fibrils with diameters of 100 nm.29 From

(a) (b) (c)6 nm15 nm50 mV

100 nm 50 nm500 nm

FIGURE 1 | Atomic force microscopy tapping mode images of (a) collagen type I fibrils on a native bone surface,37 (b) hydroxyapatite mineralcrystals extracted from bone,38 and (c) a dried film of purified osteopontin on a mica surface.39



AFM investigations of collagen type I extracted frombovine and digital tendon and rat-tail tendon, Bozecet al. further proposed to explain this preservationto be a result of a twisted arrangement of collagenmicrofibrils, i.e., like a nanoscale rope. Whereas thismodel delivers an acceptable explanation of the D-banding, it is somewhat controversial to a tube-likemechanical behavior reported earlier by Gutsmannet al.30 In their work the authors noted from AFMimaging experiments using that collagen from rat-tail tendon exhibits fibrils that seemed to have beenkinked during the preparation process. This behav-ior was found to be reversible in fibrils, which weremechanically manipulated underneath a light micro-scope, as they reversed to their original shape afterexhibiting a kink due to compression. Hence, furtherresearch is required to obtain a more complete modelexplaining both structural and mechanical propertiesof collagen type I fibrils. A basis for such modelsmight also come from AFM studies performed inforce mode, i.e., either doing force spectroscopy orindentation type experiments.Only force spectroscopyexperiments will be considered in this section, whereasindentation type investigations will be discussed in thenanoindentation part of this paper (see Section AnimalBone and Table 2). Mechanical measurements of col-lagen are important for bone research since collagencomprises about 30% of bone volume. Force spec-troscopy experiments on isolated collagen fibrils3133

have shown that these cannot be adequately describedusing worm-like chain models,3436 i.e., their behav-ior does not originate from entropic elasticity butrather from structural changes within these fibrils,that might include the irreversible rupture of bonds,such as crosslinks. In addition, Shaw et al.33 reportedan indication that the ability of collagen fibrils to dissi-pate energy during a force spectroscopy pulling eventdepends on previous sample history. In their exper-iments, fibrils that had been dried and rehydratedexhibited lower forces at rupture compared to samplesthat were kept hydrated. Despite the interesting resultsof these studies, no trustworthy quantitative materialproperties could be derived; the values of tens of mega-pascals for the elastic modulus reported by Grahamet al.32 lack explanation for how strain within the fib-ril was determined and are heavily relying on a ratherarbitrary assumption of collagen fibril radius.

Hydroxyapatite Mineral CrystalsCalcium and phosphate, in the form of carbon-ated hydroxyapatite or dahllite [CaCO3-2Ca3(PO4)2]make up the inorganic portion of bone. As mentionedabove, bone is a nanocomposite and accordingly themineral crystals have typical sizes in the range of

1300 nm. AFM is well suitable to directly ana-lyze these small mineral crystal sizes as has beenelegantly demonstrated by Eppel et al. and Tonget al. who imaged isolated mineral crystal frommature and young bovine bone (femora and tibiae),respectively.38,45 Both studies found a large numberof small crystallites with typical dimensions of 12nm thickness, 610 nm widths, and 915 nm length,as shown in Figure 1(b), which are consistent withresults from small angle X-ray scattering.76 However,Eppel et al. also reported a fraction of much largercrystals with dimensions up to 250 nm. As mineral-ization of collagen fibrils happens internally as wellas externally, they proposed that the larger crystalsmight be located in the interfibrillar space whereasthe smaller ones should be located within the gapzones of the collagen fibrils. Similarly, also a fewother AFM studies conducted on animal37,43,49 andtissue engineered40 bone have reported large crystalsizes. Hence it is not clear whether such large crys-tals as found in bovine or elephant bone would alsobe present in human bone. It is also possible thatlarge crystals are simply an artifact. Sample prepa-ration processes such as polishing or cutting couldresult in the agglomeration of individual particles intolarger particles. In any case, there is no sufficient datapresent at the moment that would allow a definitivestatement, and further research is required to clarifythe situation. Table 1 presents an overview mineralcrystal size, collagen fibril diameter, and D-banding,as determined from AFM images.

Noncollagenous ProteinsIn addition to fibrils of collagen type I and hydroxya-patite crystals, bone also consists of a small fraction(


TABLE 2 Elastic Modulus and Hardness from Nanoindentation Experiments on Human Bone Samples

Tibia, Berkovich tip, OPM, dry

Elastic Modulus(GPa)

Hardness(MPa)

Donor Notes Type Direction IndentationDepth (nm)

MaximumForce(mN)

Reference

(22.5 1.3) 614 42 2 males (57, 61 years) C (OL) L 1000 13 52(25.8 0.7) 736 34 C (IL)(22.4 1.2) 617 39 2 males (60, 66 years) C (OL) L 1000 20 53(25.7 1.0) 736 44 C (IL)(16.6 1.1) 564 34 C T(17.8 1.7) 2 females, 3 males

(6491 years)C (OL) L 500 3 54

(20.1 1.7) C (IL)(25.1 2.1) Male, 52 years Transverse is average of

indentation modulusvalues D11 and D22reported in Ref. 55

C (OL) L 700 8 55

(27.1 1.7) C (IL)(16.8 1.9) C (OL) T(19.1 1.3) C (IL)(18.6 0.9)(32.2 1.5)

Male, 52 years E varies according todifferent loading ratesand load functions

C (OL) L 600 6 56

(17.2 3.5) 422 84 Child withosteogenesisimperfecta (OI)type III

C L N/A 6 57

(14.8 1.5) 411 34 Children with OI typeIII (3.212.4 years)

C L N/A 3 58

(13.1 1.6) 413 42 TVertebra, Berkovich tip, OPM, dry (except Ref 59)


Hardness(MPa)

Donor Site Type Direction IndentationDepth (nm)

MaximumForce(mN)

Reference

13.4 2.0 468 79 2 males (57, 61 years) T-12 T T 1000 13 5222.7 3.1 664 98 1 males (61 years), 2

female (57, 66years)

L-1 T L 1000 8 60

(15.7 1.5)(18.0 2.2)

553 88572 107

T

19.4 2.3 618 61 5 male, 2 female(61 3.4 years)

T-12,L-1

T L 1000 20 53

15.0 2.5 515 82 T18.1 2.8 547 87 1 male (61 years), 2

female (57, 66years)

L-1 C L 1000 8 60

16.8 3.0 539 99 T8.8 2.28.2 1.2

310 100310 49

L-5 CT

L N/A 500 N/A 59



TABLE 2 Continued

Femur, Midshaft, OPM


Hardness(MPa)

Tip State Donor Notes Type Direction IndentationDepth (nm)

MaximumForce (mN)

Reference

19.1 5.4 234760 B dry 4 male,4 female

C (OL) N/A 500 4 61

21.2 5.3 (75.3 11.7)years

C (IL)

16.58 0.32 B dry Male, 65 years C T 1000 N/A 6223.45 0.21 L19.9 1.4 604 68 N/A dry Male, 56 years C (OL) L 200 1 6325.8 1.3 800 47 C (IL)19.5 1.5 617 64 N/A wet 10 male (4580

years)C (OL) N/A 500 N/A 59

19.5 1.5 666 87 C (IL)22.5 1.5 B dry Male, 52 years C (OL) L 150 1.5 6426.6 2.1 840 90 B dry Male, 50 years C (OL) L N/A 5 6522.9 2.7 800 130 T18.4 490 B dry Children with OI

type IIIC L N/A 3 58

19.8 510 (3.212.4 years) T

25.8 3.021.8 3.6

B dry Female 53 years In vicinity ofmicroc-rack(s) andnormal

C N/A N/A 5100 66

19.5 1.5 617 64 N/A Wet 10 male (4580years)

C (OL) L 500 N/A 59

19.5 1.5 666 87 C (IL)Femur, Neck, OPM, Berkovich tip


Hardness (MPa) State Donor Type Direction IndentationDepth (nm)

MaximumForce (mN)

Reference

(6.9 4.3)(15.9 5.1)

(234 147)(484 168)

Wet 4 female (69, 74,76, 93 years), 4male (53, 75,77, 85 years)

T N/A 500 2 67

(15.8 5.3 Dry 4 male, 4 female(75.3 11.7)years

C(OL) N/A 500 2 61

17.5 5.3 C (IL)11.4 5.6 T11.0631.6 (480 100)

(1250 250)Dry Female, 86 years C (OL) N/A 100600 N/A 68

(7.40 0.45)(18.5 4.9)

(290 130)(950 440)

Wet C (OL)

14.4 3.016.0 2.4

470 130536 89

Wet 10 male (4580years)

C (OL)C (IL)

L 500 N/A 59

10.6 2.6 380 110 T N/A



TABLE 2 Continued

Femur, Head, OPM, dry


Hardness (MPa) Tip Donor Type Direction IndentationDepth (nm)

MaximumForce (mN)

Reference

18.89 0.38 SP, B 5 donors(6785years)

Subchondral (healthy) N/A N/A 15 (SP)10 (B)

69

19.22 0.46 Articular calcifiedcartilage (healthy)

Medial

19.01 0.29 Superior18.51 0.43 SP, B 6 donors

(5589years)

Subchondral(osteoarthritic)

N/A

17.84 0.84 Articular calcifiedcartilage(osteoarthritic)

MedialSuperior

20.39 0.36 18.8 1.9 750 200 B Female,

86 yearsC L 550 5 70

22.0 2.4 1080 220 T TFemur, Other, OPM, dry

Elastic Modu-lus (GPa)

Hardness(MPa)

Tip Donor Notes Type Direction IndentationDepth (nm)

MaximumForce (mN)

Reference

12.214.2 680900 B 5 male (50, 64,65, 71 years)

Trochanter T N/A 700 6 71

21.1 1.0 N/A 2 female, 1 male(91, 89, 73years)

Proximal femur T N/A 1000 20 72

18.1 017 B Male, 65 years Condyle T N/A 1000 N/A 62ElasticModulus(GPa)

Hardness(MPa)

Method State Donor Site Type Direction IndentationDepth (nm)

MaximumForce (mN)

Reference

19.81.2 OPM Dry 2 female, 3 male(6491 years)

First metatarsalbone

C (OL) C(IL)

L 500 3 54

16.9 1.0 16.8 2.8 708 100 OPM Dry 23 female,

(64.4 5.4)years

Iliac crest T N/A 500 4 73

20.8 2.4 910 150 CSM1325 520960 OPM Dry 1 male, 2

female (54,48, 39 years)

Patella N/A 400 0.93.5 74

16.4 4.017.0 3.017.3 3.6

550 150560 110580 110

OPM Wet 10 male (4580years)

Distal radius C (OL)C (IL)C, Pri-marylamellar

L 500 N/A 59



TABLE 2 Continued

13.8 3.2 490 110 T N/A19.2 2.0 665 150 OPM Dry Children with OI

type Iii and IVIliac crest, tibia,

and femurC N/A N/A N/A 75

18.5 2.3 640 60 Iliac crest, tibia,and femur

T

B, Berkovich tip; SP, spherical tip; OPM, OliverPharr method; CSM, continuous stiffness measurement; T, trabecular (type) or transverse (direction); C,cortical; OL, osteonal lamella; IL, interstitial lamella; L, longitudinal.

the fibrils. The two most abundant of the multitudeof different NCPs are osteopontin (OPN) and bonesialoprotein (BSP). Qiu et al. have shown in an elegantin situ AFM study that OPN has a large influence onthe growth of calcium-oxalate monohydrateCalcium-Oxalate Monohydrate (COM) crystals, the majorconstituent of kidney stones.82 Habelitz et al. con-ducted similar AFM studies on amelogenin, a matrixprotein in enamel.83 Qiu et al. observed that OPNseems to selectively bind to certain COM crystal faces,hence, controlling size and shape. More importantly,networks of OPN, BSP, and dentin matrix protein1 (DMP1) have been found to have self-healing andenergy storage properties in force spectroscopy andsurface force apparatus studies.39,84,85 In these studies,purified films of each of those proteins were depositedonto a mica surface and their mechanical response totension was probed in a hydrated, close to a physio-logical, environment by pressing an AFM tip into theprotein film in the presence or absence of di- or higher-valent positive ions.When such ionswere present, highenergy dissipation (10161015 J per pull-off-event,equivalent to 1701700 covalent CC bonds) in sub-sequent pull-off events was preserved, given that thedwell time of the tip on the surface was of the order of10 s. The observed dependence of energy dissipationin the presence of positive ions and dwell time canbe explained by the fact that OPN, BSP, and DMP1are all highly negatively charged proteins, containinga high number of posttranslational phosphorylatedgroups. Hence, positive ions such as Ca2+ or La3+ canform proteinprotein as well as proteinsubstrate saltbridges, which result in an increase in energy dissi-pation and which can re-form after relaxation due tothermal motion. This is the so-called sacrificial-bond-hidden-length effect.86 Complementary investigationof OPN films with a surface force apparatus alsoshowed that the protein can serve as an energy storagewhen being compressed in the presence of Ca2+ ions.84An AFM image of a dried film of OPN is shown inFigure 1(c). The maximum thickness of this film isabout 10 nm; hence, similar thicknesses of NCP-enriched structures in bone such as cement lines couldmake this intrinsic effect of direct importance for themechanical properties of bone matrix. This is also

corroborated by earlier force spectroscopy studies inwhich two small trabecular bone samples with freshlycleaved surfaces were mounted on a substrate and anAFMcantilever.87,88 When these surfaceswere pressedtogether and then separated, similar energy dissipa-tion and molecular self-healing effects were seen asin the experiments on purified NCPs. The same effecthad already been reported earlier by Thompson et al.who performed force spectroscopy experiments onpolished surfaces of bone.89 These results suggest thatthe nanoscale energy dissipation mechanism and theself-healing character of NCPs, despite their smallamount, might be crucial for macroscopic mechani-cal properties of bone. A macromechanical proof ofthis hypothesis is missing to date. Nevertheless, it isknown that the amount of NCPs drastically decreaseswith age and disease,16 as does fracture resistanceof bone.17 Hence, there seems to be a correlationbetween the amount of NCPs in bone and its frac-ture resistance. Whether this is a causal relationship,however, remains to be elucidated.

Bone At the Nanoscale

The Mineralized Collagen Fibril: The BasicBuilding Block of BoneStudies relying on seminal work by Weiner and Wag-ner utilizing SEM have shown that at the lowest levelthese components form the building block of bone:the mineralized collagen type I fibril.19 These miner-alized fibrils were previously been thought to containhydroxyapatite mainly in their inner gap zones.18

That mineral is indeed located within these fibrilshas been shown in an elegant manner by Baloochet al. who investigated individual mineralized collagenfibrils extracted from human teeth, which they repeat-edly etched in situ with citric acid while imaging themwith an AFM.42 With each etching step, a changein collagen fibril diameter was observed along withan increasing height difference between the gap andoverlap zones of the D-banding pattern, pointing toa decrease of intrafibrillar mineral. In addition to thisinterfibrillar mineral fraction, collagen fibrils are alsocoated with hydroxyapatite crystals, which has beenconfirmed from a magnitude of studies utilizing AFM,



among other techniques. AFM images of the interfib-rillar mineral as observed on cut and fractured surfaceswere first communicated by Sasaki et al. and Knappet al.48,49 Sasaki et al. used a model of an encrustedcollagen fibril to estimate the amount of interfibrillarand intrafibrillar minerals to be 77 and 23%, respec-tively. Further, investigations showed the interfibrillarmineral crystals in significantly improved resolution,as shown in Figure 2(a),43,46,51,87,90,91 compared tothe very first reports of AFM imaging experimentson bone by Tao et al.92 Bare collagen fibrils, asobserved on native trabecular and cortical bonesurface,37,46,47,91,93,94 can also be made visible on frac-tured or cut surfaces of bone by mechanical ablationof the crystals with a water jet,46 chemical treatment(Figure 2(a,b)),48,51 or exposure to osteoclasts.44 Animportant finding from AFM imaging studies of frac-tured bone surfaces is that they generally presentsimilar morphologies: the interfibrillar mineral of themineralized collagen fibrils is exposed (Figure 2(a)).This means that bone ultimately fails in the interfib-rillar space, i.e., by delamination of the mineralizedcollagen fibrils, and in turn give rise to speculationthat a matrix situated in between these fibrils couldsignificantly influence the fracture resistance of bone.As mentioned previously, freshly fractured surfaces,similar to films of purified proteins, exhibit adhesiveand self-healing character. Hence, the hypothesis thatthe interfibrillar matrix consists of exactly this frac-tion of the organic matrix is obvious. Hints of thismatrix were first reported by Fantner et al.87 and anexample AFM image showing nonfibrillar rather thanunstructured matrix residues in between mineralizedcollagen fibrils is presented in Figure 2(c). Final proofof this hypothesis was recently achieved by stain-ing for NCPs on fracture surfaces using gold-labeledantibodies.95 A recent report by Xu and Yu claiminga layer of lipid particles beneath the mineral crystals isnot convincing, as these authors investigated bone sur-faces that were demineralized at room temperature forseveral weeks, which is expected to lead to dramaticdenaturation of the organic matrix.96

A Nanoscale Model of BoneSummarizing the results of the AFM studiesmentionedabove, it is possible to hypothesize a nanoscale modelof bone ultrastructure that takes into account the threecomponents discussed in detail: the collagen type Ifibrils, hydroxyapatite crystals, and the NCPs. Thismodel is sketched in Figure 2(d). As mentioned above,these components together form the mineralizedcollagen fibril, which contains mineral crystals in itsinner gap zones and on its surface. Since the crystalgrowth influencing ability of certain NCPs has been

shown, it is reasonable to assume that all inter- and,maybe even, the intrafibrillar hydroxyapatite crystalsare covered with a layer of NCPs. This in combinationwith results obtained from force spectroscopy andimmunohistochemistry make it reasonable to assumethat also the space in between individual mineralizedcollagen fibrils is filled with NCPs (and water).In addition, the adhesion between hydroxyapatitecrystals and collagen fibrils is most probably alsofacilitated by certain linker proteins, otherwise allconnective tissues would mineralize if Ca and P ionswere present. This protein linkage is also corroboratedby results from collagen binding and in vitrohydroxyapatite nucleation assays.77,97 Hence, thenanoscale model as depicted in Figure 2(d) includes allthese details. From this model, it seems obvious thatthe interfaces comprised of NCPs are important forstructure and mechanical function, especially whenthis construct would be loaded in tension, whereasfor loads of compressive nature, the interlocking andfriction of crystals also might play an important role.

Bone at the MicroscaleIn addition to studies with higher resolution, bonesurfaces have also been studied on the microscaleusing AFM. In many such cases, AFM was used incombination with nanoindentation studies, whichwill be discussed below, in order to localize indentlocations or select indentation area,58,68,98100 assessthe surface topography and roughness,64,71,92,101 orto investigate local morphological changes due tothe indentation process.90,102104 Other studies havefocused on the visualization of microstructural ele-ments found in bone such as resorption lacunae44,47

or osteonal lamellae.64,68,70,101

NANOMECHANICAL PROPERTIES OFBONE FROM INDENTATION FORCEMEASUREMENTSIndentation force measurements, especially in theform of nanoindentation, have become a gold stan-dard measurement to determine the elastic modulus ofbone at the tissue level in contrast to macromechan-ical measurements of the apparent elastic modulus.Nanoindentation apparatuses originally emerged inthe early 1980s,7 and have been used since the 1990sto investigate bone tissue at the micrometer and sub-micrometer range. In this section, the most commonlyused experimental and analysis methods will bediscussed. Then experimental results on both humanand animal bone will be used to discuss the correlationof material properties derived from the experimentswith parameters of the experimental protocols used.



FIGURE 2 | Fracturesurface of trabecular bovinebone exhibiting collagenfibrils coated withhydroxyapatite particles(a) and the same surfacelocation after treatment withethylenediamine tetraaceticacid (EDTA) (b), exposing thetypical D-banding pattern ofcollagen type I fibrils.51

Unstructured organic matrix(white arrows) as seen in asimilar atomic forcemicroscopy image oftrabecular bovine bone87 (c),and a nanoscale model ofbone (d).

(a)

500 nm 500 nm200 nm

(b) (c)

Untreated after EDTA treatment

Inter-fibrilar mineralIntra-fibrilar mineralCollagen fibril

nm

50200

Collagen fibril

Mineral

NCPs

(d)

Nanoindentation Methods

Force-Controlled NanoindentationTheOliverPharr MethodGenerally, most nanoindentation measurements onbone are performed in force-controlled mode, where asharp tip is driven into a sample of interest at a definedloading rate to a maximum load, which may or maynot be held for a certain time, before the sample isunloaded. From the unloading portion of the curve,the elastic modulus and the hardness of the samplecan be calculated using a method established by Oliverand Pharr,105,106 whose work is based on resultsby Sneddon.107 The so-called OliverPharr method(OPM) is to date themost commonly used formof dataanalysis for nanoindentation measurements. Typicalforce displacement graphs from nanoindentationexperiments are shown in Figure 3 along with aschematic representation of an indentation curve andthe different parameters extracted for analysis. TheOPM is based on the experimental finding that theunloading portion of indentation data can be usuallywell approximated by a power law relationship:

P = (h hf)m (1)

where P denotes load, h is the displacement, hf is thefinal depth of the contact impression after unloading,and andm are constants that depend on the indentergeometry. Indentation data are usually fitted with thisrelation by varying and m. The parameter m shouldyield values between 1.2 and 1.6 for a Berkovichindenter, which is the most commonly used geometryin studies of bone, although a few studies have reliedon a sharper cube corner tip. Using this fit, the slopeof the curve at unloading, i.e., the unloading stiffnessS can be calculated, which is related to the effective orreduced elastic modulus Eeff in the following manner:

S = dPdh

= 2Eeff

A (2)

Here is a correction factor accounting for tipgeometry, which for a Berkovich tip usually rangesfrom 1.0226 to 1.085, and A denotes the contactarea, which is a function of the contact depth hc.106

For a perfect Berkovich tip, A is given by 24.5h2c andthe contact depth itself is usually calculated from:

hc = hmax PmaxS (3)



FIGURE 3 | Typicalnanoindentation data froman experiment with52

(a) and without61 (b) holdtime at maximum load, aswell as a schematicrepresentation explainingthe analysis of theunloading part of the curveusing the OliverPharrmethod105 (c). Thevariation of the anisotropicindentation moduluscompared to the truemodulus of bone, bothcalculated from ultrasonicmeasurement data,108 isshown in (d).

00

5

10

Load

(mN)

15Tibial Osteon

(a)

(c)

(b)

(d)

Loading

Unloading

200 400Displacement (nm)

Loading

Pmax

M, = 90M, = 0E, = 90E, = 0

x2-x3 plane

x1-x3 plane

10

12

14El

astic

mod

ulus

, E (G

Pa)

Inde

ntat

ion

mod

ulus

, M (G

Pa)

16

18

20

22

1001

0

1Load

[mN] 2

3

4

0 100 200Displacement [nm]

300 400 500 600

0 15 30 45Angle from x1-x2-plane, (degrees)

60 75 90

x2-x3 plane

x1-x3 plane

hmaxDisplacement, h

hf

SUnloading

600 800 1000

Load

, P

Cortical bone

Trabecular bone

with = 0.75 being the generally used value.106

Substituting Pmax and S from Eq. (1) and usingh = hmax, the relationship between hc and hf can beshown as:

hc = hmax hmax hfm (4)

Finally, the material properties of elastic modulus Eand hardnessH are determined by using Eqs (5 and 6):

1Eeff

= 1 2

E+ 1 i

2

Ei(5)

H = PmaxA

(6)

where denotes Poissons ratio and the suffix idenotes the material constants of the indenter tip.For practice this method relies on a good calibrationof the load frame compliance as well as the tiparea function. Whereas the load frame complianceparameter is more critical for large indentations in

samples with high elastic modulus, the area functionis a more critical calibration that will affect allmeasurements. In any case, both parameters can easilybe assessed using a calibration material with knownelastic modulus and performing several indentationexperiments with different contact depths, fromwhichthe needed information can then be obtained througha simple fitting procedure.105 It is important to keepin mind that the OPM is based on a number ofassumptions,61,105,108 which are as follows:

1. The constitutive behavior of the sample iselastic with time-independent plasticity withoutdamage.

2. The equations describing the elastic unloadingof a flat semi-infinite half-space are the same asfor an indented surface.

3. The Poissons ratio of the sample is known.4. The sample is an isotropic solid

When considering these assumptions, it is obvi-ous that probably only assumption number 2 holds fora nanoindentation experiment on bone samples. All



other assumptions are surely not met, which explainspart of the variation in reported values for elastic mod-ulus and hardness. Another part of the variation canbe explained by the heterogeneity of bone, resulting inlocal difference in composition and multiscale organi-zation of the bone volume element tested. Hence, giventhe need of characterization technique, it is essential tocounteract the individual influences by careful designof the experimental protocol as well as being aware ofthe limitations of the technique during interpretationof acquired data. Addressing the assumptions one byone, the following statements can be put forward:

1. Bone is not an elastic material with time-independent plasticity. In fact, bone plasticitymight depend critically on time, especially intension. However, the effect of faster transientsof bone plasticity can be reduced in nanoinden-tation experiments by using a few indentationcycles prior to the actual test cycle or byholding the peak load during the indentationmeasurement prior to unloading. Long holdingtimes might however influence the hardnessmeasurement due to creep, hence, Fan and Rhoproposed to combine several pre-cycles and atest cycle with a loading time on the order of 30s for best results.56 These authors also investi-gated the effect of loading rate for which theycould derive power law relationships, similar tomacroscopic mechanical testing experiments,109

pointing out the viscoelastic character of bone.Similar investigations by Donnelly et al. also ledto the conclusion that varying maximum loadcan significantly influence material parametersas deduced by the use of the OPM;101 however,a similar investigation by Mittra et al. did notshow any significant differences in hardnessor elastic modulus arising from changes inmaximum load or loading rate.110 It should benoticed thatMittra et al. used significantly lowerpeak loads (


imposed by bone, quantitative values must be treatedwith a grain of salt. It should be noted that this is muchless of an issue for comparative studies, which aim topoint out differences among various sample groupswith all samples prepared and tested in a similar man-ner. As can be seen from Tables 2 and 3, there is quitea variability in material constants with the used proto-col, which has also been shown systematically by Fanand Rho as well as Mittra et al.56,110 as mentionedabove. Hence, parameters obtained by the OPM withthe exact same protocol might be very well compa-rable but cross-comparison among different studiesemploying different protocols is not trivial. To over-come this issue, some work has been done on refiningand expanding the OPM and expanding it to mea-sure plastic and viscoelastic response of bone. Oliverand Pharr themselves mention a few improvementsand corrections of their original technique,106 suchas correction for pile-up, which is not encountered inbone, however.116 More interestingly, they propose tocompute the quotient of maximum load and squaredunloading stiffness for different indentation depths.This combined parameter is independent of the areafunction (Eqs 2 and 6) and should be a material con-stant; hence it can be used to check validity of theOPM. In addition Oyen has pointed out that it is rela-tively easy to assess the amount of plastic deformationin bone in addition to the original hardness param-eter of the OPM.117 More recently Tang et al. haveexpanded the stiffness parameter in Eq. (2) with aterm accounting for viscoelasticity allowing to deductthe elastic rather than apparent unloading stiffness,and they have also put forward an expression thatallows to calculate the viscosity from protocols thatlet the sample creep prior to unloading.118 It has yet tobe shown whether plasticity and viscoelasticity exten-sions can unify the spread of quantitative parameterspresented in the literature to date, but these develop-ments are promising and surely with the improvingexperimental protocols, the technique should arrive ata stage where derived parameters become reproducibleeven when using different protocols.

Other MethodsBesides using the force-controlled indentation withina single loadingunloading cycle, there are a few otherexperimental techniques that have been applied tobone, delivering similar and complementary data tothe ones delivered by the OPM.

Already mentioned by Oliver and Pharr is thecontinuous stiffness measurement (CSM) in whicha force or displacement oscillation (1100 Hz)is superimposed on a normal test protocol.106

Advantages of the technique are that it delivers

data constantly when loading or unloading thesample, and its capability (using phase information) todetermine both storage (elasticity) and loss modulus(damping/viscoelasticity). The technique itself hasbeen used for bone to limited extent.66,73,134 In a studybyWang et al., CSM shifted values for elastic modulusand hardness to higher values but overall trendsamong different groups of samples were consistentcompared to values derived using the OPM.

Balooch et al. have proposed a modulus map-ping (MM) method combining a nanoindenter withan AFM xy stage to map material properties over asurface.145 MM, similar to CSM, relies on an oscil-lating tip, which in the case of MM is scanned overa surface at constant load preset on which the forceoscillation is superimposed. Similarly to CSM, loss andstorage maps of a sample can be deducted. These canbe compared to topography images of the same loca-tion retrieved by using the indenter tip in AFM mode.MM has been used in few studies to date,98,100,122 andit seems that MM yields results quite comparable tothe ones retrieved with the OPM at similar indentationdepth.100

Very recently a scratch testing approach has beenproposed byWang et al., who were able to show a dif-ference in scratch depth and energy per scratch volumecomparing bones frommiddle-aged donors to the onesfrom elderly ones.146 In addition Mullins et al. haveproposed the use of a cube corner tip and rather highloads (up to 3N) to create cracks during indentation inorder to measure fracture toughness. Both techniquescould be important additions to the field, but theirvalue can only be determined with further studies.

Experimental ResultsThis section will briefly discuss the most importantexperimental results reported for human and animalbones and list extracted parameters of elastic modulusand hardness in two tables.

Human BoneHuman bone in its healthy state and the changes itis undergoing during age and disease are of greatinterest to the research community, as mentionedin the introductory part of this review. Quite a fewstudies on human bone have been published sincethe original article on the OPM in 1992, dealingwith changes due to aging and disease as mentionedbut to a large part also with quantifying anisotropyand heterogeneity of bone matrix material properties.The individual findings of all studies on human bonethat were considered for this review are summarizedin Table 2. Comparing the different studies, it turns



out that the reported values vary a fair amount, e.g.,the elastic modulus values of trabecular bone fromdifferent sites vary from 6.9 to 23.5 GPa, whereasthe ones for cortical bone vary from 7.4 to 31.6GPa. Essentially, this mirrors the individual results inqualitative manner; there is no clear trend for corticalbone having a higher elastic modulus than trabecularbone. In fact six of the studies reported higher elasticmodulus for cortical bone,52,53,58,61,63,75 whereasthree studies reported higher elastic modulus for tra-becular bone.60,62,70 In addition both bone types havebeen investigated in terms of their anisotropy andall results considered show that bone has generallya higher elastic modulus in its longitudinal comparedto its transverse direction.53,55,58,60,65 Moreover, fivestudies report that the elastic modulus of interstitiallamellae in cortical bone is higher compared toosteonal lamellae52,53,55,61,63 in contrast to a singlestudy claiming the opposite.54 AFM images showingosteonal human lamellae, with and without indents,are presented in Figure 4(a) and (b). Figure 4(c) showsdata from a study by Rho et al., depicting the decreasein elastic modulus of individual lamellae within anosteon when going from the Haversian channel inthe middle to the outside of the osteon.63 The elasticmodulus of human bone has also been shown to beproportional to the mineral content of the tissue74 andcompressive regions in cortical bone have been foundto have increased elastic modulus.54 It has also beenfound that sample preparation and state influencesextracted parameters of elastic modulus and hardness.Hengsberger et al. have reported significantly differentvalues for dry and wet samples.68 These authors alsoshowed that dried samples change their propertieswithin 1 day after preparation; however, after thistime no more change could be detected.70 Comparingelastic moduli determined from nanoindentationexperiments to acoustic microscopy or ultrasonic

measurements yields good correlation.55,62 Finally, astudy by Wang et al. surprisingly found no significantdifferences in iliac crest biopsies obtained from twogroups of donors: healthy individuals and ones thathad experienced vertebral osteoporotic fractures.73 Itis reasonable, however, that changes in bone qualitywill manifest themselves differently among the varioussites within the skeleton. Hence, it is possible that theiliac crest is not a predictor of material properties ofvertebral bone.

Animal BoneIn addition to human bone, animal bone also is ofinterest to the research community for a number ofreasons. First of all, animal bone is easier to obtainthan human bone; secondly, animal studies are widelyused to determine the influence of genotype on bonematrix material properties,100,102,118,119,124129 toinvestigate the efficacy of drugs,98,122,133 to investi-gate changes upon ovariectomy-induced osteoporosisand/or diet,99,122,130 and the influence of implants orfracture healing processes.119,121,132 The individualfindings of all studies on animal bone that wereconsidered for this review are summarized in Table 3.Overall, these studies showed that among the differentspecies, the elastic modulus of trabecular and corticalbone varied from 2.4 to 33 GPa and from 2 to 45.8GPa, respectively. Interestingly, all studies consideredreported in overall agreement that the elastic modulusis higher for cortical bone than for trabecularbone,126,133 that in cortical the elastic modulus ishigher in the longitudinal compared to the transversedirection,135,138,140142,144 and that the modulus ofinterstitial lamellae is higher compared to the one ofosteonal lamellae.136,140 Since discussing all of thesefindings in detail is beyond the scope of this review,a few selected findings will be briefly discussed. Forfurther details the interested reader is referred to

2420.0

10.0

020.0 m10.00

0

25.0

50.0 200.0 nm

100.0 nm

0.0 nm

50.025.00m

(c)(b)(a)22

20

18

16

14

Elas

tic m

odul

us [G

Pa]

12

100 5

y = 0.4072x + 21.775R2 = 0.82

Nr. of Lamella (from center)10

FIGURE 4 | Atomic force microscopy images of a bone lamella without64 (a) and with101 (b) indentations. The change in elastic modulus ofindividual lamellae when from the center to the periphery of an osteon is shown in (c), data from Ref. 63.



the respective original publications. Of the extensivework on the influence of genotype on elastic modulusand hardness of bone, the communication by Baloochet al. is probably the most exciting one to date. Theseauthors identified that the TGF- signaling pathwayplays an important role in specifying bone matrixmaterial properties.100 Essentially, they found thatoverexpression of TGF- leads to a decrease of elasticmodulus, hardness, and mineralization. Similarly, Taiet al. and Roy et al. reported that elastic moduluschanges accordingly with mineral concentration.90,141

In addition Roy et al. also showed a correlationbetween osteocalcin content and elastic modulus,which might support findings from studies on theimportance of NCPs for the mechanical propertiesof bone mentioned earlier. Interestingly,the elasticmodulus obtained from nanoindentation does not nec-essarily correlate with other mechanical tests or largersamples. Although Hengsberger et al. reported good

agreement between nanoindentation and microme-chanical traction experiments,135 Silva et al. foundthat comparing the elastic moduli of two differentmouse strains led to completely reversed results whenusing nanoindentation or three-point bending.125 Itshould be noted here, though, that Hengsberger et al.essentially performed a test at one hierarchical level oforganization above the level of nanoindentation andalso took into account the tissue porosity, whereasSilva et al. performed a whole bone test involvingmany more hierarchical levels of organization. Hence,the disagreement between nanoindentation and three-point bending could very well be due to changes atother levels of organization. Two recent investigationsthat were pushing the technique to its limits used AFMcantilevers for indentation experiments. In an elegantstudy,Wenger et al. used AFM cantilevers indentationof isolated collagen fibrils in a dried state and reportedvalues for the elastic modulus of 3.7 to 11.5 GPa.104 In

TABLE 3 Elastic Modulus and Hardness from Nanoindentation Experiments on Animal Bone Samples

Porcine Bone, OPM


Hardness(MPa)

Tip State Breed Site Type IndentationDepth (nm)

MaximumForce(mN)

Reference

7.78 0.47 189 15 N/A Wet Sinclair miniswine,2 years

Alveolar bone, 1 monthafter implantation ofTiO2 screw, areaclose to implant(150 m)

C

7.415.68.017.0

B Dry 8 pigs (standardDutchcommercialhybrid race), fetaland newbornanimals

Mandibular condyle T 1000 8 120

New Zealand White Rabbit, OPM, dry


Hardness(MPa)

Tip Site Type IndentationDepth (nm)

Maximum Force(mN)

Reference

2.4 0.5 N/A Spine (between L-5 andL-6) fused with aHA/-TCP scaffold, 7weeks post-surgery

T 2000 25 121

26.6 2.3 990 140 B Left distal femoral condyle T, lamellar N/A 0.5 10120.3 2.1 950 110 T, interlamellar20.6 1.6 830 60 T, lamellar 319.9 1.2 820 90 T, interlamellar



TABLE 3 Continued

Murine (Mus) Bone, OPM (except Refs 118,122,123), dry


Hardness Tip Genotype/Age Site Type Direction IndentationDepth (nm)

MaximumForce

Reference

28.8 5.8 1.00 0.36GPa

N/A C57B6/DB. 3 months Proximalfemur

C N/A 500 3 mN 124

29.2 5.2 1.08 0.42GPa

COX-1/, 3 months

28.8 4.2 1.12 0.13GPa

COX-2/, 3 months

29.5 1.3 1.11 0.09GPa

B SAMR1 (wild-typecontrols), male andfemale, 4 months

Femur C L 500 6 mN 125

31.4 2.7 1.16 0.12GPa

12 months

27.8 1.8 0.980.09GPa

4 months Tibia

30.8 1.8 1.11 0.09GPa

12 months

32.1 1.6 1.23 0.12GPa

SAMP6, male andfemale, 4 months

Femur

33.7 2.5 1.31 0.13GPa

12 months

29.9 2.1 1.11 0.13GPa

4 months Tibia

33.1 1.3 1.26 0.08GPa

12 months

26.6 0.6 1070 30MPa

B C57BL/6J (B6) female, 4months

Femurmidshaft

C L N/A 6 mN 126

13.6 1.2 690 20 MPa Proximalfemur

T N/A

28.3 1.3 1220 30MPa

Tibiamidshaft

C L

28.5 0.5 1110 20MPa

DBA/2J (D2) female, 4months

Femurmidshaft

C L

14.71.3 730 30 MPa Proximalfemur

T N/A

22.9 2.6 1410 30MPa

Tibiamidshaft

C L

30.5 0.8 1080 30MPa

C3H/HeJ (C3) female, 4months

Femurmidshaft

C L

15.3 1.6 830 30 MPa Proximalfemur

T N/A

30.9 1.6 1460 20MPa

Tibiamidshaft

C L

45.8 1.0 2.64 0.07GPa

B SW129 wild-type,


TABLE 3 Continued

17.41 0.51

627 35 MPa D5 mutant, 8 weeks

20.49 0.68

774 30 MPa Wild-type (C57Bl6,C57Bl6/129, D6D2),8 weeks

27.32 0.68

887 35 MPa DNTRII mutant, 8weeks

29.02 0.68

928 47 MPa Smad3 +/ mutant, 8weeks

27.0 1.5 887 41 MPa Smad3 / mutant, 8weeks

4.60 0.42 B BALB/cByJ, female, 1day

Tibia midshaft C N/A N/A 1 mN 128

4.80 2.2 2 days7.60 1.3 7 days11.90 3.7 14 days20.6 5.7 21 days23.7 2.2 30 days29.3 3.6 40 days33.9 450 days

28.3 1.325.9 1.6

1220 3 MPa829 58 MPa

B C57L/6J, female, 16weeks

Tibia midshaft,fresh and afterstorage inETOH

C L N/A 6 mN 129

30.9 1.528.0 1.6

1460 18MPa880 76 MPa

C3H/HeJ, female, 16weeks

Tibia midshaft,fresh and afterstorage inETOH

C L

14.2 2.6 592 172MPa

B C57Bl/6N Forelimb (OPMw. viscositycorrection)

C N/A N/A 10 mN 118

22.1 0.7 872 60 MPa ICR24.2 5.7 CC Swiss Webster, Male

and female, 6.7months

Vertebra (L-5),(ModulusMapping)

T N/A N/A 0.51 N 122

25.0 3.4)24.1 3.6 1160 260

MPaB Female Bg-Nu-Xid

mice, 2 yearsCalvarial bone,

(Doerner Nixmethod)

C N/A N/A 0.15 mN 123

27.3 10.5 1.0 0.7 GPa B C3H/HeN mice, 1014weeks

Femur C N/A 100, 200,500

0.5, 1, 7 mN 103

23.2 2.3 784 94 MPa B C57BL/6J mice, 16weeks

Femur stored inethanol

C L N/A 0.2 mN 57

Murine (Rattus) Bone, dry

Elastic Modulus Hardness (MPa) Tip Method Breed/genotype Site Type IndentationDepth

MaximumForce (mN)

Reference

16.1 3.9GPa

670 239 B OPM SpragueDawley,female, 37 weeks;sham @ 17 weeks

Vertebra(L-4)

T 500nm N/A 130



TABLE 3 Continued

15.8 3.9GPa

690 260 SpragueDawley, female,37 weeks;ovariectomized @17weeks

Vertebra (L-4) T

22.39 0.84GPa

823 45 N/A OPM Wild-type (2 weeks) Tibia C N/A 6 131

19.9 2.1GPa

795 21 Femur

(15.1 1.3)(21.8 1.6)GPa

(600 73)(1021 61)

N/A OPM SpragueDawley, female,1.5 years; sham @ 6months

Vertebra (L-5),anterior,posterior andlateralcortical bone

C 900 nm N/A 99

(12.3 2.2)(21.3 2.4)GPa

(500 110)(876 29)

SpragueDawley, female,1.5 years; ovariectomized@ 6 months and on lowprotein diet

(13.82.5)(20.5 1.2)GPa

(626 87)(967 85)

SpragueDawley, female,1.5 years; ovariectomized@ 6 months, on lowprotein diet, partiallyrescued with essentialamino acid supplement

0.61 1010MPa

50 mcono-sphericaltip

OPM SpragueDawley, male, 5months

Femur, fracturecallus(granulationtissue,chondroidtissue, andwoven bone),35 days afterfracture,

C 13 m 0.020.3132

20.18 0.36GPa19.35 0.39GPa

854 19849 21

B OPM Fischer, female (110 weeks)normal diet

Vertebra (L-5) CT

900 nm N/A 133

19.62 0.36GPa19.33 0.40GPa

921 17888 21

Fischer, female 110 weeks,strontium ranelate (900mg/kg/day) diet

CT

12.92 0.42GPa12.37 0.34GPa

484 23457 18

Fischer, female 110 weeks,normal diet

CT

14.59 0.30GPa14.24 0.37GPa

458 18510 19

Fischer, female 110 weeks,strontium ranelate (900mg/kg/day) diet

CT

5.011.5 GPa Atomic forcemicroscopy(AFM) softtappingcantilever

OPM N/A Tail tendon,individualcollagenfibrils

N/A 510%of fibrildiame-ter

N/A 104

24 4 GPa 920 180 N/A N/A SpragueDawley, 10months

Femur T N/A 0.3 98



TABLE 3 Continued

Canine Bone, Berkovich Tip, dry


Hardness (MPa) Method Breed Site Type IndentationDepth (nm)

Maximum Force(mN)

Reference

13.0 4.2 530 240 N/A 5 mature dogs Distal femora T N/A 0.250.75 11016.8 1.419.5 1.312.5 1.0

771 53888 59

OPMCSMFEA

Beagle dogs Lumbar vertebra T 500200500

4N/A

134

Bovine Bone


Hardness(MPa)

Tip Method State Site Type Direction IndentationDepth (nm)

MaximumForce

Reference

27.2 4.2 N/A OPM Dry Femurmidshaft

C L 900 N/A 135

21.0 2.6 T(24.4 2.2)/(21.1 2.0)

(578 52) /(680 102)

B OPM dry/wet Femur C (OL) L 1000 20 mN 136

(27.5 1.2) /(25.1 1.6)

(818 49) /(730 48)

C (IL) L

12.9 2.9 B OPM Dry Tibia C N/A 300 1mN 909.00 0.35 FEA13.424.2 B OPM Dry Femur

midshaftC L 9001000 7mN 137

25.4 5.4 B OPM Dry Femurmidshaft

C N/A N/A 5100 mN 66

(11.8 3.6)(14.1 5.3)

AFM stifftappingcan-tilever

OPM Dry Proximaltibia

C L 50 5 N 138

(8.5 3.7)(14.6 5.0)

T

16.6 2.9 844 159 B OPM Dry Tibia C (Plexi-form)

L N/A 0.3 mN 57

Other Species, dry (except Ref 139)

ElasticModulus (GPa)

Hardness(MPa)

Tip Method Species Site Type Direction IndentationDepth (nm)

MaximumForce(mN)

Reference

40.2 5.165.3 6.1

B OPM Whale (Balaenopteraphysalus)

Whale (Mesoplodondensirostris)

Tympanic bulla(fin)

Rostrum

C N/A N/A 5100 66

20.4 N/A OPM Equine(thoroughbread), 2years

Tibia C (OL) L 600 4 140

13.5 T

22.6 C (IL) L

14.2 T

10.9 1.1 N/A OPM,CSM

Fallow deer (Damadama)

Antler, dry T N/A 2000 N/A 139

5.4 0.9 Antler, wet15.9 1.2 359

62N/A N/A Carp (Cyprinus

Carpio)Rib (1st to 5th

pair)C L 1000 N/A 141



TABLE 3 Continued

8.86 0.74 249 25 T(19.4 2.2)(3.8 0.7)

B OPM Herring (Clupeaharengus)

Intramuscular;fullymineralized toearlymineralizedregion

N/A L 700 8 142

(9.1 1.6)(3.5 0.3)

T

(8.1 1.4)(1.1 0.4)

(676 156)(84 20)

B OPM Wild-typeZebrafish

Vertebra N/A N/A N/A 5 102

(14.7 1.5)(6.8 1.9)

(699 270)(304 139)

Stopseldtl28d

Zebrafish

9.0 2.1 573 119 N/A OPM Wild-type (Tu)Zebrafish

Vertebra N/A N/A 1000 10 143

6.1 2.4 367 177 Mutant (lil/lil)Zebrafish

18.7 1.2 730 47 B OPM(correctedaccordingto Ref.136)

Sheep (OvisAries)

Femur midshaft C L N/A 5 144

13.3 1.6 546 62 TB, Berkovich tip; CC, cube corner tip; OPM, OliverPharr method; CSM, continuous stiffness measurement; MM, modulus mapping technique; FEA, finiteelement analysis; T, trabecular (type) or transverse (direction); C, cortical; OL, osteonal lamella; IL, interstitial lamella; L, longitudinal

FIGURE 5 | Residual indentation in a bovine cortical bone90(a) and a collagen fibril from a rat-tail tendon104 (b).

(a)200 nm

400 nm Cant

ileve

r

(b)

Figure 5 an AFM image of an indent in a collagen fibriland as a comparison an indent in bovine are presented.Tai et al. used AFM cantilever-based indentation tocreate elastic modulus maps, which were fed into finiteelement simulations of a three-point bending as well asan indentation test.138 In both cases, the nanoscale het-erogeneity led to an improved energy dissipation andductility.

Future TrendsNovel developments in the realm of indentationtechnology might soon make this technique availableto the clinician for minimally invasive measurementson patients;147,148 however, it remains to be shown

whether a few local indentation measurements can aidthe diagnosis of bone health and fracture risk. Also,trends to investigate the higher harmonic oscillationsof AFM cantilevers might be of use to create nanoscalemodulus maps in a matter of a few minutes.149

CONCLUSION

The results discussed in this review show how AFMand nanoindentation have led to further our under-standing of the structurefunction relationships ofbone.Whereas AFMhasmostly been essential to studystructure and nanoindentation to study themechanicalproperties of bone, the techniques might merge more



in the future with novel technological developments.The nanoindentation experiments discussed show thatthe elastic modulus and hardness as well as viscosityand plastic deformation can be deducted from them.Although there are still several issues with the tech-nique itself when applied to bone, it has become themethod of choice, especially in mouse models, fordetermining bone matrix material properties, such aselastic modulus and hardness. Anisotropy and hetero-geneity of bone surely pose some issues and leave roomfor needed improvements of the technique. In con-trast existing expansions to the standard OliverPharrmethod accounting for time-dependent plasticity andviscoelasticity need to be evaluated in more studies

and should essentially be introduced as a standardanalysis. Despite the fact that this will to some extentcomplicate data analysis, these additions could lead toan improved stability of derived elastic modulus andhardness values for different experimental protocols.Obviously, dynamic techniques such as the CSM,which inherently separates elastic from viscoelasticcontributions, might also improve the situation. Inany case, when planning standard nanoindentationexperiments using the OPM, it is advisable to consultthe literature and use protocols employed by otherauthors on similar bones or research questions in orderto be able to compare results to previously reportedvalues.

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Atomic force microscopy and indentation force measurement of bone