In connective tissue, the role of collagen fibers is to:

11.11 Conclusions

Collagen fibers are the structural elements found in vertebrate tissues that transmit forces, store, and dissipate energy. Collagen fibers limit the deformation of tendon and other load bearing tissues and have a hierarchical structure that includes collagen molecules, microfibrils, fibrils, fibers, and fascicles. Collagen molecules are packed into a quarter-stagger arrangement with neighboring molecules staggered by multiples of D, which is about 22% of the molecular length. During mechanical deformation collagen molecules as well as the gap region of the D period are stretched. At larger strains, molecules and fibrils slide by each other, which leads to energy losses. Finally, collagen fiber failure occurs by disintegration of some of the hierarchical structure yielding collagen subfibrils that lose much of their mechanical strengths.

The ability of collagen-cell interactions to provide dynamic structural alterations in the mechanical properties of ECMs provides clinicians with the ability to monitor changes in tissue structure. However, this will require new techniques to measure changes in properties that occur during the disease process. For this to occur a detailed understanding of collagen biomechanics is necessary since the time-dependence and deformation mechanism must be considered in order to provide accurate interpretations of the mechanical properties. This also requires the use of “gold standards” to make sure any new test developed can duplicate the behaviors that have been reported in the literature in the pioneering research of Yamada (1970) and Fung (1973).

Future studies are needed to identify the changes that occur in collagen fibers and their mechanical properties found in tissues that are associated with cancers and other pathologies (Dudea et al., 2013). Collagen fibers in ECMs are oriented in more than one direction and form multilayered sheets such as is observed in skin, cartilage, and bone. Changes in packing patterns and fiber orientation may be a useful criterion for early diagnosis of a number of disease states.

8.6.7.5.1 The Collagen Fiber Attachment Stretch λfa

Collagen fibers are in a continual state of deposition and degradation in the current configuration κ. Vascular cells (fibroblasts in the adventitia and vascular smooth muscle cells in the media) work on the collagen fibers to attach them to the matrix in a state of stretch in this configuration. Consequently, the recruitment configuration κR is inferred from the stretch at which fibers are configured and attached to the extracellular matrix in the current configuration κ. On this note, it is important to recognize that the matrix is pulsating and thus it is desirable for the definition of the attachment stretch to explicitly take this into account. Watton et al.[315] hypothesized that the fibers are configured to the matrix to achieve a maximum stretch during the cardiac cycle and introduced the terminology attachment stretch where λfa≥1 .

4.3.2.1 Tension.

Collagen fibres are the main tension resistant elements in connective tissues. Their presence in articular cartilage suggests that tensile stresses are present, even though the tissue is loaded predominantly in compression perpendicular to the articular surface.

Investigators have examined the tensile properties of thin slices of articular cartilage (Kempson 1979; Bader et al. 1981; Bader 1985). When an isolated dumbell-shaped specimen of cartilage is subjected to a tensile force in a plane parallel to the articular surface, the resulting stress-strain behaviour is non-linear in form, as illustrated in the curves in Figure 4.5.

The response can be regarded as a continuous progression of three phases of behaviour. Initially at low levels of stress the collagen fibres tend to become aligned in the direction of the tensile force. The tensile tangent modules in this phase depends on the initial orientation of the collagen fibres and the effective resistance of the proteoglycan gel to their alignment. With increasing alignment of the collagen fibres the component of the applied stress along the fibres increases and the stress versus strain curve increasingly reflects the mechanical properties of the fibres. With increasing stress, therefore, the tangent modulus increases until fracture of the specimen occurs within the gauge region.

There is some variation in the tensile properties with orientation of the specimen. Superficial layer specimens have been shown to highly strain-limiting in tension, thereby implying least compliance, along the split-line direction and least strain-limiting across it. However, despite earlier reports to the contrary, a recent study has revealed no strong dependence of fracture strength on split-line orientation (Kamalanathan and Broom 1993). Tensile properties also decrease with respect to distance from the articular surface. These results demonstrate both the anisotropic and non-homogeneous nature of articular cartilage.

The dependence on the relationship between the two main structural components and tensile properties can be demonstrated by the incubation of cartilage specimens in proteolytic enzymes (Bader et al. 1981; Bader 1985). These studies demonstrated that the effect of proteoglycan degradation alone was limited to a reduction in the tensile stiffness in the initial phase of the stress-strain relationship, as seen with curve b in Figure 4.5. However, damage to the collagen network with associated release of proteoglycans from the matrix was sufficient to reduce both the tensile stiffness at all levels of stress and the tensile strength of cartilage, as shown with curve c in Figure 4.5. Indeed this response was evident with one specific enzyme, leukocyte elastase, despite no measurable release of collagen from the matrix. These results confirmed the importance of the covalent intermolecular cross-links, which elastase are known to disrupt (Starkey et al. 1977), on the overall mechanical properties of the collagen fibrils in articular cartilage.

In two related tensile fatigue studies by Weightman and colleagues (1976; 1978), it was demonstrated that human articular cartilage is prone to tensile fatigue failure in vitro. Tests on specimens from human femoral head cartilage revealed that the fatigue resistance decreased with age at a rate, which was faster than that which could be explained by normal usage. The decrease in fatigue resistance could not be related to either of the two major solid constituents of the specimens. By combining the observations the following model was produced:

(4.4)S=23−0.1a−1.83logN

where S is the stress in MPa, a is the age in years and N is the number of cycles to failure.

In a recent in vitro study, the effects of partial fatigue on the tensile properties of articular cartilage were evaluated (McCormack and Mansour 1998). Cartilage specimens were repetitively loaded at a frequency of 0.75 Hz using a maximum compressive load of 65 N for approximately 80,000 cycles. This procedure revealed no macroscopic damage to the surface of the cartilage specimen. Subsequent tensile tests produced a reduction in tensile strength, which was attributed a weakening of the interfibril connections which link collagen fibrils in the middle and deep zones of the cartilage matrix (Broom 1984).

1 Introduction

Collagen fibers form the extracellular framework of all tissues and are essential for tensile strength. The collagen molecules in these fibers are synthesized as larger precursor proteins, known as procollagens, from which large propeptide domains are enzymatically cleaved en bloc. As such, the synthesis rate of fibrillar collagens can be determined by measuring procollagen propeptides in blood. Assays are available for the two most common human fibrillar collagens, type I and III [1,2].

Because collagens are normal tissue constituents that undergo constant renewal, procollagen propeptides are not tumor markers per se. In contrast, malignant tumor growth needs to invade surrounding tissue via degradation and new synthesis of the extracellular matrix. In soft tissues, this process can be a continuous fibroproliferative reaction. Although resembling wound healing, this phenomenon is not self-limiting in cancer [3,4]. Procollagen propeptides released from a fibroproliferative reaction of malignancy can potentially be used as markers of tumor behavior.

Bone, a special case of connective tissue, practically determines the concentration of circulating type I procollagen propeptides due to its relatively high turnover rate. In cancer, however, these concentrations can be affected by therapy as well as development of skeletal metastases [5,6].

This review summarizes the biochemistry and physiology of procollagens and their propeptides. The fibroproliferative process in response to cancer and bone turnover is presented. Interaction of cancer cells with the extracellular matrix and the process of bone metastases are also addressed.

3.3.3.1 Organisation of the bone matrix

This classification criterion is mainly based in the orientation of the collagen fibres in the bone matrix. The three bone matrices, which will be defined, are related at the same time by other classification criteria such as the vascularisation, the shape of the bone cells or the degree of mineralisation, among others.

2 Microstructural Properties: Osteons and Trabeculas

Mineralized collagen fibers form into planar arrangements called lamellas (3–7 μm wide). In some cases these sheets (lamellas) of mineralized collagen fibers wrap in concentric layers (3–8 lamellas) around a central canal to form what is known as an osteon. The osteon looks like a cylinder about 200–250 μm in diameter running roughly parallel to the long axis of the bone. Other forms of cortical bone, where the mineralized collagen fibers are less well registered and no pattern can be distinguished, are called woven bone. In some forms of bone, the lamellas are overall tangential to the outer surface of the bone (without forming osteons), and together with woven bone tissue, form a larger plywood-type stacking of thicker layers (150–300 μm) around the complete circumference of the bone. Cancellous bone is made of an interconnecting framework of trabeculas in a number of combinations all comprising of the following basic cellular structures: rod–rod, rod–plate, or plate–plate. A trabecular rod is about 50–300 μm in diameter.

Over the years Ascenzi and co-workers have examined the mechanical properties of single Haversian systems in bending (Ascenzi et al. 1990), tension (Ascenzi and Bonucci 1967), compression (Ascenzi and Bonucci 1968), and torsion (Ascenzi et al. 1994; Table 2). Osteons with longitudinal lamellas are better for tension and torsion and perhaps stronger in bending as well, while osteons with alternating lamellas are more suited for compression. The stiffnesses produced from torsion testing of individual osteons were much higher than those for the whole bone. Ascenzi et al. do not justify or explain this result. However, this effect may be explained by the use of Cosserat micropolar elasticity, which allows for local rotation at cement sheaths and other internal boundaries (Lakes 1995). Katz (1981) confirmed that there is a majority of fibers in the longitudinal direction in osteons and also along the long axis of trabeculas in cancellous bone, but the situation in general is muddled. Marotti (1993) claims that fibers in general follow two patterns which constitute thin and thick lamellas; the thin ones are more oriented and compact, the thick ones are more diverse and sparse (somewhat microporous) in their elements.

Table 2. Mechanical properties of single osteons and trabeculas. Longitudinal (L) osteon has marked longitudinal spiral course of fiber bundles in successive lamellas. Alternate (A) osteon has fiber bundles in one lamella making an angle of nearly 90° with the bier bundles of the next one.

a Ascenzi et al. 1990, b Ascenzi and Bonucci 1967, c Ascenzi and Bonucci 1968, d Ascenzi et al. 1982.

The trabecular properties are easier to study in isolation. However, in spite of recent reports (Choi et al. 1990, Rho et al. 1993) there remains some controversy regarding the value of the elastic modulus of single trabeculas (Table 3). Trabecular bone material properties are important for characterizing various bone pathologies, and the remodeled bone adjacent to various joint implants, because they are affected by disease sooner than cortical bone. In the past it was assumed that individual trabeculas, single osteons, and a thin cortical shell possessed the same mechanical properties as those of large cortical bone specimens regardless of their type or size (Wolff 1892). However, many investigators produced values for the elastic modulus of individual trabeculas, single osteons, and a thin cortical shell that were considerably less than that for whole bone (Choi et al. 1990, Rho et al. 1993, Ashman and Rho 1988). Mechanical testing of cortical osteonal samples has shown a similar modulus, while individual trabeculas show a lower modulus of elasticity than that of large cortical bone specimens. This issue needs further clarification.

Table 3. Mechanical properties of single trabeculas.

a Wolff 1892, b Runkle and Pugh 1975, c Townsend et al. 1975, d Williams and Lewis 1982, e Ashman and Rho 1988, f Ryan and Williams 1989, g Hodgskinson et al. 1989, h Kuhn et al. 1989, i Mente and Lewis 1989, j Choi et al. 1990, k Rho et al. 1993, l Rho et al. 1997.

A literature survey of measured and estimated values of the modulus of trabecular bone material shows that moduli values range from 1 GPa to 20 GPa (Table 3). Rho et al. (1993) showed that the relationship derived from this data, between elastic moduli and density in cancellous bone material, could not be extrapolated from similar data from tests on cortical bone material and its density. This led the authors to conclude that the materials of the two bones were intrinsically different. However, the data are not definitive, as the dependence of the modulus on other parameters such as location, microstructure, or density variation was not considered. The most recent studies by use of nanoindentation (Rho et al. 1999) and by finite element analysis (FEA) simulation (vanRietbergen et al. 1996) suggest that in fact the elastic properties of single trabeculas are very similar to the properties of nearby cortical tissue. This matter requires further clarification because imprecise values may lead to misinterpretations of the structural function of each type of bone material or misinterpretation of the role of trabecular bone in the mechanical behavior of normal and implanted joints.

Tapered Fibres

The collagen fibres in sea cucumber dermis and sea urchin ligament are tapered rather than cylindrical. The shape is ensured by the nucleation and growth mechanism (Trotter et al., 1998; Trotter et al., 2000a) by which the fibres are formed (Fig. 8). We will see that this shape reduces the volume of fibre that is not exploited at close-to-maximum load-bearing capacity. Therefore, supramolecular self-assembly in this instance has the further advantage of ensuring that the tensile strength of the fibre is exploited efficiently.

In the case of conventional, cylindrical reinforcing fibres, the diameter is the same at all points along the fibre length. If, as is expected, the fibres deform less readily than the matrix, the shear stress in the matrix at the fibre–matrix interface is largest at the fibre ends (see Kelly and Macmillan, 1986, equation 6.49), while the tensile stress in the fibre is greatest at the middle of the fibre (see Kelly and Macmillan, 1986, equation 6.40). These results are summarised in Fig. 9. Ideally, discontinuous fibres will be long enough for the stress at their midpoint to approach the fibre failure strength. The necessary length depends on the fibre radius, in a manner that is easily and commonly (Kelly and Macmillan, 1986) quantified as follows. With reference to Fig. 10, consider a small length dx of a fibre, near one end. The tensile stress in the fibre increases by dσ over this distance. This increase in tensile stress is achieved by the interfacial shear stress τ acting on the interface area accommodated within the distance dx. A simple force balance

(3)τ(2πrdx)=dσ(π r2)

can be rearranged to give

If the failure stress σuf of the fibres is reached over a transfer distance xuf, Eq. 4 can be integrated between the corresponding limits to obtain an expression that incorporates this distance:

Explicit integration of Eq. 5, followed by rearrangement, leads to:

So xuf is an increasing function of r: an increased fibre thickness will increase the capacity of the interfacial surface area to transfer load to the fibre (in proportion to r), but will also (and even more effectively) increase the capacity of the fibre cross-section to carry that load (in proportion to r2).

It follows that much of the material in the fibre is wasted, in that the tensile strength is not being properly exploited along almost the whole length of the fibre! This problem is exacerbated in thicker fibres. The load near the ends of the fibres could be carried adequately by a thinner fibre cross-section, compared to the load near the middle of fibres. A less wasteful use of material, and a more efficient exploitation of the fibre properties, would therefore be achieved if the fibres were to taper from the middle towards the ends. Also, regions of the fibre having a smaller cross-section would then be able to undergo a larger elastic deformation, thus matching more closely the deformation of the matrix; therefore, the shear stress concentration in the matrix near the fibre ends would be reduced. These gains have been recognised and discussed qualitatively in the context of sea cucumber and sea urchin collagen fibres, which are appropriately tapered (Trotter and Koob, 1989; Trotter et al., 1994, 2000b).

It is instructive to consider how the force balance in Eq. 3 and the critical half-length (transfer distance) of the fibres in Eq. 6 are affected by allowing the fibres to taper towards their ends. Reference should be made to Fig. 11, which again considers the shear stress acting on a small length dx near the end of a fibre. The tensile stress in the fibre again increases by dσ over this distance, and the fibre radius increases by dr. From the geometry of Fig. 11,

(7)−drdx=tanθ;i.e.r=xtanθ

and the interface area accommodated in the distance dx is

(8)Area=2πrdx+212πdrdx=2πrdx+πdrdx≈2πrdx

provided that θ is small so that the second-order term containing the product of two differentials can be ignored. θ is indeed found to be small for the collagen fibres under consideration here: their length-to-width ratio is of the order of 2000 (Trotter et al., 2000b).

A simple force balance now gives

(9)τ(2π rdx)cosθ≈τ(2πrdx)=dσ(πr2)

Substitution of Eq. 7 into Eq. 9, followed by simplification, leads to

If the stress in a fibre increases from a very small value σif (at a point xif that is arbitrarily close to the end of the fibre) to the failure stress σuf (at a point xuf that ideally is at the midpoint of the fibre), Eq. 10 can be integrated between the corresponding limits to obtain an expression in which xuf again represents the transfer distance

(11)∫σ ifσufdσ=2τtanθ∫ xifxufdxx

Explicit integration of Eq. 11, followed by rearrangement, then leads to:

(12)σuf−σif=2τtanθIn xufxif

Therefore

(13)σuf≈ 2τtanθlnxuf−C

where C is a constant that subsumes events at the end of the fibre. The adoption of a small, non-zero lower limit when performing the integration is a device that circumvents the need to consider the logarithm of zero in the calculation. A similar approach is conventionally used when considering the strain energy associated with a screw dislocation in a crystal (Cottrell, 1953): a small volume around the dislocation line is assigned a ‘core energy’ that is not calculated explicitly.

Eq. 13 can be rearranged to give

(14)xuf=exp[tanθ2τ(σ uf+C)]

So xuf is an increasing function of tanθ: the fibre is shortest, promoting the most effective use of reinforcing material, if the taper is gradual. This is precisely what sea cucumbers and sea urchins do.

4.4.1.3 Corresponding Permeability

The presence of collagen fibers restricts the permeation of interstitial fluid within cartilage, leading to a correspondingly higher diffusivity in directions parallel to local fiber orientations. Considering the collagen fibers, we determine the filtration velocity nF wFS as (cf. Pierce et al., 2013a,b).

(4.66)nFwFS=KF(−gradp+ρFRb),

where KF is a positive definite material parameter tensor representing the intrinsic permeability of the cartilage solid matrix and b is the body force per unit mass. We propose KF as (Ricken and Bluhm, 2010; Pierce et al., 2013b, 2016)

(4.67)KF=k0S(nF1−n0SS)mH,H=14π∫Ωρ(a0)I4(a)a⊗adΩ,

where H is a spatial structure tensor defined by the integration of ρ(a0) over all normalized, spatial fiber orientations aˆ=λ−1FSa0 (with λ=|a|), k0S [m4/Ns] is the initial Darcy permeability, and m is a dimensionless parameter controlling the general isotropic deformation dependence of the permeability (cf. Eipper, 1998). Inclusion of the volume fraction nF relates to the change of permeability caused by the change of pore space, where n0SS denotes the reference solid volume fraction. Here we used H, via ρ(a0), to define the range of permeabilities resulting from ideal alignment of collagen fibers to an isotropic distribution of the collagen fibers.

11.2 Intervertebral disk components and structure

Water, collagen fibers and proteoglycans molecules represent 90 to 95% of IVD components [BIB 01]. The rest is composed of minor molecules and cells maintaining and renewing the IVD tissues. With 6000 cell/mm3, the IVD has a really low cell density compared to other joints cartilages where it is usually around 14000 to 15000 cell/mm3 [BIB 01, MAR 75, RAN 00, RAN 04]. Within the IVD the cell distribution is related to cell nutrients and metabolic wastes routes. Being avascular the IVD cell’s environment is renewed thanks to blood capillaries surrounding the annulus fibrosus and blood marrow through vertebral endplates [HOL 81, RAN 04, SHI 10]. The cell population is mainly composed of:

fibroblasts responsible for collagen type I production especially within outer part of the annulus fibrosus [BIB 01];

chondrocytes that produce collagen type II and proteoglycans in the whole disk.

These cells are responsible for the IVD tissue homeostasis-producing extracellular matrix of which the predominant components are collagen, elastin and proteoglycan molecules [BIB 01, KRA 09].

11.2.2 Intervertebral disk annulus fibrosus

The annulus fibrosus is made of 7 to 15 concentric layers clearly visible at macroscopic scale as presented on Figure 11.3. Usually lamellae tend to be thinner in posterior regions than in anterior ones. This tissue is the link in between two adjacent vertebrae with one to two thirds of the lamellae fixed to the vertebral endplates [MAG 09, SHI 10]. The annulus fibrosus is assimilated to a hyaline cartilage with 60 to 85% of water content [ACA 95, RAN 04]. As a fiber-oriented cartilage, from inner to outer IVD the fiber orientation decreases in regards to the circumferential direction (see Figure 11.2) from 45° to 25° with an average value around 28° [AMB 09, MAL 12, NOA 05, SHI 10].

Figure 11.3. Macroscopic pictures of the annulus fibrosus oriented fiber structure

adapted from [BAL 13]

Made of collagen type I, this structure gives an anisotropic behavior to this tissue as a composite material. Being soft, these fibers only play a role undergoing tensile loadings [AMB 09, BAL 13].

19.2.1.1 PLA and copolymers

For example, collagen fibers were incorporated onto PLA fibers to mimic blood vessel properties [19]. PLA and copolymers are widely used as scaffolds due to their biodegradability. These polymers are hydrophobic and have good mechanical strengths, also the hydrolytic product is lactic acid which is catabolized in the lactic acid cycle into water and CO2 [80]. Stitzel et al. developed small diameter vascular grafts based on a single helical wind of collagen fibers embedded in a PLA nanofiber matrix, to mimic the stress environment similar to that found in native arteries [19]. The construct was seeded with human aortic SMCs and after 10 days culture, they demonstrated that SMCs proliferate under static conditions. Comparing their results with a scaffold without collagen fibers, the authors observed that cells were oriented along the principal stress line, demonstrating that they can influence the cell alignment.

Jeong et al. fabricated novel tubular scaffolds, made of a porous jellyfish collagen tube blended with poly(lactide-co-glycolide (PLGA) fibers, using freeze drying and electrospinning processes for vascular TE [20]. PLGA is a hydrophobic, biocompatible, and biodegradable polymer and is a FDA approved polymer for human use. This polymer is widely used in drug delivery and TE applications. These new scaffolds had good mechanical properties, were durable under the mechano-active system, and possessed tissue compatibility to SMCs and ECs. The authors cocultured SMCs and ECs under a pulsatile perfusion system. They observed that pulsatile flow influenced in cell alignment and promoted cell proliferation and retention of differentiated cell phenotype if compared with static culture condition. Hu et al. [11] prepared microtubular scaffolds by thermal-induced phase separation (TIPS) technique using PLGA (70/30) modified by plasma treatment and then with collagen anchorage method. Cell affinity of the new scaffold was tested with A10 cells in vitro and the authors observed a high viability and cells grew well and fast. Stitzel et al. [21] engineered vascular graft by electrospinning method using collagen, elastin and PLGA (50/50) to improve their mechanical properties and were tested in vitro with bovine endothelial and smooth cells. The authors demonstrated that the new vascular scaffolds promote proliferation and maturation as well as support long-term cell growth. On the other hand, Stitzel et al. demonstrated the biocompatibility of the scaffolds in vivo, thus the author implanted subcutaneously the scaffold in mice. After 4 weeks, the authors observed that the novel scaffold did not cause toxicity or inflammatory reactions, demonstrating that can be used as a potential vascular graft.