Resumo Direito civil

Journnl qf Orthopnedir Resecir(.h 4~379-392, Raven Press, New York 0 1986 Orthopaedic Research Society

Tensile Properties of Human Knee Joint Cartilage: I. Influence of Ionic Conditions, Weight Bearing, and

Fibrillation on the Tensile Modulus

Shaw Akizuki, *Van C. Mow, ?Francisco Muller, $Julio C. Pita, §David S. Howell, and ?Daniel H. Manicourt

Biomechanics Research Laboratory and *Department of Mechanical Engineering, Aeronairtical Engineering and Mechanics, Rensselaer Polytechnic Institute, Troy, New York; and +Internal Medicine Department, $Department of

Medicine, and §Arthritis Division, University of Miami, Miami, Florida, U.S.A.

Summary: The flow-independent (intrinsic) tensile modulus of the extracel- Mar matrix of human knee joint cartilage has been measured for normal, fibrillated, and osteoarthritic (removed from total knee joint replacements) cartilage. The modulus was determined in our isometric tensile apparatus and measured at equilibrium. We found a linear equilibrium stress-strain behavior up to -15% strain. The modulus was measured for tissues from the high and low weight-bearing areas of the joint surfaces, the medial femoral condyle and lateral patello femoral groove, and from different zones (surface, subsurface, middle, and middle-deep) within the tissue. For all specimens, the intrinsic tensile modulus was always less than 30 MPa. Tissues from low weight-bearing areas (LWA) are stiffer than those from high weight-bearing areas (HWA). The tensile modulus of the ECM correlates strongly with the collagen/proteoglycan ratio; it is higher for LWA than for HWA. Osteoarthritic cartilage from total knee replacement procedures has a tensile stiffness less than 2 MPa. Key Words: Intrinsic tensile modulus-High and low weight bearing areas- Normal, fibrillated, and osteoarthritic cartilage- Ion concentration effects- Correlation with biochemical composition.

Articular cartilage exhibits both viscoelastic and biphasic responses under tension, compression, and shear (18,31,32,Sl). Even under simple uniaxial tension, it exhibits mechanical behaviors that de- pend on strain rate, ionic conditions in the bathing fluid, biochemical composition, and the structural organization of the dominant components of the tissue (collagen fibrils and proteoglycans)

Address correspondence and reprint requests to Dr. V. C . Mow at Department of Orthopedic Surgery, Columbia-Presby- terian Medical Center, 622 W. 168th St., PH5-129, New York, NY 10032, U.S.A.

Dr. Akizuki’s present address is Department of Orthopaedic Surgery, Shinshu University, Medical School, Matsumoto, Japan.

(15,23,25,36,40,.50). Thus, to understand the influence of composition and structure on the mechanical behavior of normal and degenerating human articular cartilage, all these factors must be consid- ered and carefully controlled. In this investigation we assessed the influence of all these factors on the equilibrium or intrinsic tensile modulus of normal, fibrillated, and osteoarthritic knee joint cartilage.

To explain the motivation for our experimental protocol, a brief description of the structural organization of the extracellular matrix (ECM) of articular cartilage is required. The main components of the ECM are collagen fibrils, proteoglycans, and water. The collagen fibrils are organized into a fine meshwork of definite architecture, with interfi-

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380 S. AKIZUKI ET AL.

brillar spaces filled with proteoglycans and water. It is generally believed that the elaborate structural features of proteoglycan aggregates are essential for the organization of the ECM by promoting physical, chemical, and mechanical interactions and/or entanglement sites with the surrounding collagen network (17,34). Together, the collagen network and the proteoglycans contained within the network form a cohesive fiber-reinforced com- posite matrix with all the essential characteristics of a porous-permeable solid material filled with water (32). This porous-permeable solid matrix provides the ability for the tissue to withstand the enormous stresses of joint articulation (1,9,10,20) in situ under conditions of large deformation (3).

Each of the structural components of the ECM is endowed with specific characteristics that provide the ECM with some outstanding mechanical properties. Native collagen fibrils are strong in tension (6,21,40,50), and proteoglycans are highly hydro- phylic (41). In addition, proteoglycans exhibit an elastic behavior when compressed (16). In the ECM, negatively charged proteoglycans are con- fined to about 20% of their free-solution volume (17). Thus, a large Donnan osmotic swelling pressure is exerted by the trapped proteoglycans onto the surrounding collagen network. This swelling pressure is balanced within the ECM by the tension developed in the constraining collagen network (28).

The internal equilibration of forces may be dis- turbed chemically by changing the counter-ion concentration in the external bathing solution. An increase of the Na+ concentration in the bathing solution will cause a decrease of the Donnan osmotic swelling pressure of the proteoglycans (27), thus decreasing the tensile stress acting in the collagen network. Myers and co-workers (37) have shown that for isometrically stretched cartilage specimens, the measured change of tension associated with a change of ionic conditions in the bathing solution may be theoretically modeled and predicted. Thus, one objective of this investigation was to use this method to assess the modulation of the tensile stiffness of the ECM of human knee joint cartilage by changing Na+ concentration.

COMPOSITIONAL AND STRUCTURAL HETEROGENEITIES

The composition of cartilage varies significantly over the joint surface (46). These variations appear

to be related to joint loading (7,11,39,45). In general, regions that are habitually highly loaded have more proteoglycan content than regions that are habitually lightly loaded, and regions that are habitually lightly loaded have more collagen content than regions that are habitually highly loaded. This variation provides a convenient way to study the effect of natural variation of biochemical composition on the tensile modulus of the ECM. In addition, layerwise variation of collagen ultrastructure, composition, and structural anisotropies for this tissue a re well known (5,12,33). Thus, tensile samples must be carefully controlled in terms of the depth from within the tissue. In this investigation, specimens were taken from four distinctively different ultrastructural zones (surface, subsurface, middle, and middle-deep). All specimens were taken from an orientation, as close as practically possible, “parallel” to the local split-line direction (23,4030). Where possible, tissues from fibrillated joint surfaces and some osteoarthritic surfaces were tested to assess the effect of these degenera- tive processes.

LOAD CARRIAGE IN CARTILAGE

Load carriage within articular cartilage is shared between the interstitial fluid and the intrinsic material properties of the ECM. For example, during rapid compressions, the frictional drag associated with interstitial fluid flow may account for, de- pending on the rate of compression, more than 90% of the compressive stiffness; the stiffness of the ECM contributes to the remainder of the measured compressive stress (24). This is a consequence of the nonlinear biphasic behavior of the tissue (32). Thus, in order to determine the “intrinsic” flow- independent properties of the ECM in compression, equilibrium measurements must be made or measurements must be made under infinitesimal shearing conditions (18) to avoid a false high value for the intrinsic moduli of the ECM resulting from this flow-generated stiffening effect.

The same situation applies for determining the intrinsic tensile properties of the ECM. Constant strain-rate tensile experiments (23,40,50) yield tensile stiffnesses that a re not flow-independent (25,26). For tensile specimens 1-2 cm long, an imposed displacement rate of 4-5 mm/min will cause a significant flow-generated stiffening effect (25). For example, for human medial femoral condyle cartilage surface zone, Kempson (22) determined a

J Orthop Res, Vol. 4 , No. 4, 1986

TENSILE MODULUS OF HUMAN KNEE CARTILAGE 381

tensile modulus ranging from 40 MPa (for a specimen obtained from a donor more than 85 years old) to 125 MPa (for a 20-year-old specimen) when the applied stress was 5 MPa. In similar constant strain-rate experiments on 2-year-old bovine cartilage, both Woo and co-workers (50) and Roth and Mow (40) found tensile moduli to be less than 30 MPa when the applied stress was 5 MPa. This discrepancy may be due to differences in the composition and structure between bovine and human cartilage, ages of the specimens, and fibrillation, or it could be due to the mechanical testing protocol. In this investigation, to avoid ambiguous interpreta- tions and the possibility of the flow-generated stiffening effect, all tensile moduli of the ECM were measured at equilibrium.

In addition, Woo and co-workers (51) showed that similar cartilage specimens exhibit pronounced stress-relaxation effects in tension and that this behavior may be described by the quasi-linear viscoelastic theory proposed by Fung (14). Hence, it is difficult to interpret correctly the results of constant strain-rate tensile experiments where both the flow-generated stiffening and stress-relaxation effects are occurring simultaneously. Once again, these difficulties may be circumvented if equilibrium measurements are made.

Thus, in this investigation the intrinsic tensile modulus of the ECM were determined at equilibrium. The variation of this property with the naturally occurring differences of tissue composition and surface fibrillation were assessed. Specimens were categorized as being from high and low weight-bearing areas (HWA and LWA) and from different zones throughout the depth of the tissue (surface, subsurface or transition, middle, and middle-deep), and whether they were normal, fibrillated, or osteroarthritic (OA). These intrinsic tensile moduli of the ECM were measured in deionized water and 0.15 M NaCl solution.

MATERIALS AND METHODS

Specimens and Tissues

A total of nine knee joints were tested (Table 1). Six knees were obtained from the Tissue Bank and Surgical Research Laboratories' at the University

I Courtesy of Dr. Theodore I. Malinin, Director, Tissue Bank and Surgical Research Laboratories, Department of Surgery, University of Miami.

TABLE 1. Characteristics of the human knee joint specimens

Joint no. Age Sex Side Condition Criteria

1 24 Male 2 31 Male 3 35 Male 4 42 Male 5 51 Male 6 60 Male 7 65 Female 8 68 Female 9 74 Female

Left Right Left Right Right Right Left Right Right

Whole Whole Whole Whole Whole Whole TKR parts TKR parts TKR parts

Normal Normal Normal Normal Fibrillated Fibrillated Osteoarthritic Osteoarthritic Osteoarthritic

TKR, total knee replacement.

of Miami. These joints were divided into two groups: group 1 -joints with visually normal cartilage with no India staining over the surface (ages: 24, 32, 35, 42 years) and group 2-joints with par- tial fibrillation of the cartilage surface visible to the naked eye and stained with India ink (age: 52 and 60 years). In this group, India ink stain grades of two and three, according to the scheme developed by Emery and Meachim (13,29), were noted over some aspects of the joint surface. However, most of the specimen surfaces did not exhibit any India ink uptake. Parts from three osteoarthritic joints were obtained from total knee replacements (TKR) (ages: 65, 68, and 74 years)2. In this study, we define group 1 specimens as normal, group 2 specimens as fibrillated, and TKR specimens as OA.

Following the protocol established at the Tissue Bank and Surgical Research Laboratories, knees harvested from fresh cadavers were stripped of all soft tissues and flash-frozen in liquid nitrogen. The specimens were then placed in heavy-duty plastic bags, sealed, packed with dry ice, and shipped in an insulated box by overnight delivery to Rensse- laer. All specimens were received frozen and were placed immediately in a freezer and stored at -20°C until ready for dissection. Parts from the TKRs were received from the operating room after sufficient amounts were taken for ordinary patho- logical examinations. They were placed in plastic bags, frozen, and sent; all specimens arrived frozen and were subsequently processed in the same manner.

Preparation of Tensile Specimens

In this study, cartilage from the distal femur was

These joints were obtained from Queen's University Hos- pital, courtesy of Dr. Derek T. V. Cooke, and Women's and Brigham Hospital, courtesy of Dr. Clement B. Sledge.

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382 S . AKIZUKI ET AL.

used. Care was taken to harvest and categorize specimens from what is generally regarded as the HWA and LWA regions over the femoral condyles (9,10,43). The lateral patellofemoral groove (LPG) and the medial femoral condyle (MFC) sites were chosen because their broad surfaces make it easy to obtain tensile specimens (Fig. 1).

Our methods and procedures for preparation of tensile specimens have been reported (37,40), but to briefly summarize, rectangular osteochondral blocks were cut from the LPG and MFC. These were then divided into several blocks, some from the HWA region and others in the LWA region. The exact locations of the HWA and LWA regions for each joint, of course, are not known. Thus, the HWA and LWA classifications should be regarded only as a qualitative guide. Cartilage was then removed from the subchondral bone at the tidemark and subsequently sectioned using a freezing stage Leitz sledge microtome. The superficial 50 km of the tissue was microtomed from each block of cartilage to yield a flat surface. Subsequently, seven to ten slices, -250 krn thick, were removed serially from each block (Fig. 2). In some blocks, due to the thinness of the tissue or fibrillation, specimens from the middle-deep and deep zones could not be obtained, which explains the unequal number of samples in our data pool (Table 2). From each slice of tissue, two to three 1.8 x 15 mm tensile strips may be obtained. The strips were taken, as close as practically possible, “parallel” to the local split- line direction (21,40). However, in some cases, particularly with the LPGiHWA specimens, a slight angle may exist between the axis of the specimen and the split-line direction.

In total, we successfully tested 130 strips of normal cartilage from joints showing no signs of India ink staining, 47 strips of “fibrillated cartilage’’ (obtained immediately adjacent to regions stained with India ink), and 23 strips of “os-

teoarthritic cartilage” (obtained adjacent to severely osteoarthritic lesions or regions with no cartilage). In severely osteoarthritic TKR materials, only LWA specimens were obtained because no HWA cartilage existed. The number of specimens in each anatomical category are indicated in Ta- ble 2.

Biochemical Analyses

Initially, we prepared -100 mg of tissue for each category of specimen for the complete set of biochemical assays. To obtain this amount of mass, the remaining material in each slice and the immediately adjacent slices (above and below) were grouped together to yield the required mass. The mechanically tested strip was not included in the -100 mg of tissue for biochemical assay. In fibrillated and osteoarthritic cartilage, some of the materials pooled in these samples necessarily included tissues from regions of frank fibrillation, which un- doubtedly must influence the biochemical results. With all the samples, we measured hydroxyproline, hexosamine, glucosamine, galactosamine, uronic acid, collagen, proteoglycan, and water contents.

Water content was determined and expressed as the ratio of the wet weight minus the dry weight relative to the wet weight of the specimen. Prior to weighing, each sample harvested for biochemical analyses was equilibrated in 0.15 M NaCl solution for 1 h. Upon removal from the solution, extra fluid covering the surfaces of the specimen was gently removed using soft absorbent tissue paper, and the specimen was weighed immediately in a micro- balance (with an accuracy of 0.01 mg) to determine its wet weight. The specimen was then dried in a vacuum oven at 60°C for 48 h. The dry weight was then determined and the water content calculated.

Dried cartilage specimens were subjected to pa- pain digestion with an enzyme-to-substrate ratio of

Lateral Patello Femoral Groove ( L P G )

FIG. 1. Schematic diagram of location of high weight-bearing area (HWA) and low weight-bearing area (LWA) on the distal femoral condyle (right). Five sample locations were chosen to be in areas of relative high and low weight-bearing regions over the joint surface (left). The exact nature of loading over a specific joint is, of course, not known

ow Weight Bearing

H,qhWe,ght

Lateral Femoral Medial Femoral Condyle (LFC) Condyle (MFCI

J Orthop Res, Vol. 4 , No. 4 , 1986


Depth prn

Surface Zone ( S ) Sub Surface Zone ( S S)

-Middle Zone ( M 1

-Middle Deep Zone ( M D) 2000

Subchondral Bone I Split Line Direction

FIG. 2. Osteochondral block of material removed from the femoral condyle was sliced into 250 pm thick slices. Usually, eight to ten slices could be harvested from each block of material. The top 50 pm was removed to obtain a flat planar surface. Each slice was subsequently cut into strips such that the long axis of the strip was aligned parallel to the local split-line direction. The length and width, L and w, of the strip equal 15 and 1.8 mm, respectively.

1: 100 (wtlwt). After digestion, the material was centrifuged at 10,000 rpm for 10 min, the superna- tant was removed and filtered, and known volumes of sample solutions were made. Total hexuronate was determined in an aliquot of the solution using the carbazole method as modified by Heinegard (19). The remaining solution was divided into two parts that were independently hydrolyzed in HCl 6 M at 100°C for 17 h, and then evaporated to dry- ness at 60°C under a carefully controlled circulation of dry air. One of the residues was dissolved in 1 .O ml of buffer solution containing 0.25 M citric acid, 0.60 M sodium acetate, 0.80 M NaOH, and 0.20 M acetic acid. Hydroxyproline was determined in this solution by the method of Woessner (49) using hydroxyproline gold labeled standards from Aldrich. The second residue was dissolved in the alkaline sodium acetate buffer as described by Blu- menkrantz and Asboe-Hansen (8) and modified by Wagner (47). This method was used to determine the total hexosamine and the galactosamine to glucosamine ratio. In these determinations, galactosamine HC1 (Sigma G-0500) was used as standard.

Collagen was calculated by assuming hydroxy-

m m - 0

J Orthop RPS, Vol. 4, N O , 4 , 1986


proline residues in 100 FW, and proteoglycan content was calculated by assuming 22% hexuronate in the monomer. Initially, we restricted the sample size to 100 mg for the biochemical analyses for each category of tissue, e.g., LPG/HWA, because we wished to correlate the site-specific biomechanical properties with the site-specific biochemical composition. However, we found this to yield samples that were too small for accurate biochemical assays. Thus, the same zones of different categories, e.g., surface zones of the LPG/HWA and MFClHWA, were grouped together for these biochemical assays. Consequently, in our statistical correlations, mean values of biomechanical properties of each zone were correlated with the corresponding biochemical data.

Dimensional Measurements of Cartilage Strips

In order to calculate the engineering stress (force/original cross-sectional area) acting on the specimens, an accurate cross-sectional area must be determined. Further, as both normal and osteoarthritic tissues swell upon a change of salt concentration in the external bathing solution (fibrillated and osteoarthritic cartilage swelling much more than normal cartilage), the dimensions of the tensile specimens were determined after complete equilibration in deionized water and again after complete equilibration in 0.15 M NaCl solution.

The thickness was measured in a new apparatus designed to sense the electrical conductivity of the tissue. The tensile strips were gently placed on a precision-ground stainless-steel plate attached to a Leitz electro-optical micrometer. The tip of the micrometer, also electrically connected to the circuit, was aligned perpendicular to the ground steel plate and at a known distance above the plate. The ver- tical traverse of the tip was controlled by a precision electric motor. The downward traverse of the tip was stopped when the tip touched the electrically conducting tissue. The distance traveled by the tip yielded a reading for the specimen thickness. The accuracy of the devise has been verified to be within 23.0 pm against an optical noncon- tacting (but much more tedious) method. Thickness measurements of the specimens were made at ten points along the length of the strip and averaged. All measurements were completed within 1 min to minimize evaporation effects.

The lengths and widths were measured at ten and five points along the respective edges, using a

Bausch and Lomb stereo-zoom microscope adapted with an Olympus precision X-Y translation stage coupled with a Microcode digital micrometer, and averaged. The difference in the length of the strip equilibrated in deionized water and 0.15 M NaCl solution must be taken into account in the calculation of the true tensile strain for the 0.15 M NaCl equilibrated strip (37).

Tensile Experiment

Figure 3A shows our isometric tensile apparatus (ITA) used to determine the equilibrium tensile properties of the ECM. [This apparatus is also especially effective in studying the kinetic swelling behavior of cartilage strips held in isometric tension (37). This phenomenon, for human articular cartilage, was measured and is discussed in a com- panion paper, see Akizuki et al. (2).] The principle of the apparatus is very simple: elongation of the specimen is determined by the movement of an air- activated piston (Fig. 3A). The displacement of the piston, and hence the right jaw, is controlled by a micrometer stop. The micrometer provides a simple method for determining the exact jaw-to-jaw length of the specimen. In this way, the precise amount of stretch experienced by the specimen may be i re determined.^ The load is measured by the load cell attached to the stationary left jaw. Hence, the entire load history, including the equilibrium load, is easily monitored. In addition, pro- visions for a stepwise change of the bathing solution are provided by the continuous stream of fresh solutions from two separate sources attached to two rows of shower nozzles (Fig. 3B). Thus, the tensile properties of the specimen equilibrated in various solutions may be determined in a single experiment.

Figure 4 provides the complete details of the experimental protocol for the determination of the equilibrium tensile properties of cartilage strips as well as the kinetic swelling properties (2). Figure 4A shows the sequences for displacement and sa-

Recent results have indicated inaccuracies in using jaw-to- jaw strain measurements (40,51) and inhomogeneities in the strain field in other soft tissues in similar tensile experiments (52). We have nevertheless used this measure of strain and assumed it to be uniform in our calculations for the tensile modulus because of the extreme difficulties encountered in measuring gage-length strains with our Optron electro-optical noncon- tacting extensometer when a continuous stream of fluid bathes the specimen.

J Orlhop Res, Vol . 4, No. 4 , 1986


3A

OMETER STOP

PISTON ROD

MOVING J A W

//

38

F

F

linity changes imposed on the specimen. After mounting the specimen in the ITA, preconditioning, and equilibrating in deionized water, a very small amount of stretch was imposed onto the specimen at time T,. This stretch caused an immediate load response, followed by stress relaxation occurring during T,-T, (Fig. 4B) . Complete relaxation usually occurred within 15 min (37,51). The equilibrium stress attained during the TI-T, period was adjusted to be below 1 gf. We used this protocol to define the tare loads and tare lengths of the specimens. This tare load was chosen because it is the load required to maintain an initially “straight” configuration for the naturally curled strips. The tare load was subtracted from all subsequent values

TER

FIG. 3A and B. Schematic diagram of the isometric tensile apparatus (ITA) used in the experiments. See text for a description of its function. Figure 3B is reproduced with per- mission from Myers et al. Journal of Biomechanical Engi- neering, ASME, 1984.

of force in our calculations for the stress, and the tare length was used as the initial length in our calculations for strain when the specimen was equilibrated in deionized water. Calculations of strains for specimens equilibrated in 0.15 M NaCl were adjusted by the amount of free contraction between deionized water and 0.15 M NaCl determined under free swelling conditions.

The kinetic isometric swelling (2) and the tensile experiments began at time T, (Fig. 4). At this time, the bathing solution was changed stepwise to 0.15 M NaCl from deionized water (Fig. 4A), and the resulting time variation of stress change, i.e., the nonequilibrium swelling kinetics, was monitored during T,-T, (Fig. 4B). Upon equilibration in 0.15

J Orthop Res, Vol. 4, No. 4, 1986


PROTOCOL OF EXPERIMENT

- Displacement Chonge .... NaCl Cancentrotion Change

C 0

t 5G

T i m e

( b ) FIG. 4. Experimental protocol used to determine intrinsic tensile moduli in different bathing solutions. The solid line shows the displacement history applied and the dashed line shows the history of change of the bathing solution imposed on the specimen (a). Typical changes in stress due to the applied strain and imposed change of salinity are shown (b). See text for an explanation of the stress history.

M NaCl and at T,, deionized water was reimposed on the specimen in a stepwise manner; the resulting kinetics of stress change was monitored during T,-T,. Following reequilibration in deionized water and at T,, an additional increment of strain was added onto the specimen by the ITA in a stepwise manner (Fig. 4A); the resulting stress relaxation was monitored during T,-TS. Upon equilibration (at T,) the NaCl solution was applied and the resulting kinetics of stress change again monitored. This pro- cedure of cyclical changes of bathing solution and increments of strain was repeated eight times, until 15% strain was reached. No attempts have been made by us, and none have ever been reported in the literature, to determine whether articular cartilage, when stretched up to 15%, will return to the tare length upon removal of the tensile stress. We believe, however, that at the 15% strain level, collagen alignment within the ECM might be irrevers- ible. Further investigations are required to quantita- tively assess the residual tensile strain upon re-

moval of tensile stress. In compression, it is well known that articular cartilage does fully recover upon removal of the compressive stress (44).

Shown in Fig. 4B is the typical mechanical stress-relaxation pattern after each increment of strain and the typical kinetic swelling behaviors of the cartilage specimens at each level of strain (37). Note that the nature of stress change associated with the imposition of the 0.15 M NaCl solution de- pends on the tensile strain; at low strains, T,-T,, a contraction effect occurred, while at high strains, T,-T,, an expansion effect occurred. At interme- diate strain levels, T,-T,, a transition occurred between the contraction effect and the expansion effect. This pattern existed for both bovine and human articular cartilage. For this investigation, the equilibrium tensile moduli of the ECM in deionized water and in 0.15 M NaCl were determined by simply calculating the ratios of the increments of the equilibrium stress, i.e., points D, E, F and points A, B , C of Fig. 4B, to the increments of strain (Fig. 4A), respectively. In the range of strains chosen (- 15%) the equilibrium stress-strain response of the ECM in tension was remarkably linear (Fig. 5). Figure 5 shows some typical results for a subsurface zone specimen bathed in deionized water and 0.15 M NaCl. We note that, over the years, we have repeatedly studied, for control pur- poses, the effect of freezing on the biomechanical properties of bovine, porcine, and dog articular cartilage. For storage at -20°C up to as long as 1 year, no consistent pattern of biomechanical property changes were noted.

Tensile Modulus in Delanued water

I p'I Tensde modulus in 0 15M Nacl 1' /

0 a E v, / v, 0.50 / W [L t- v,

0.25 ,I//

0 I I I I I - 0 0.0 5 0.10 0.1 5

STRAIN FIG. 5. Equilibrium tensile stress-strain relationship of the extracellular matrix for a subsurface specimen. The relationship, under the present experimental protocol, is linear up to -15% tensile strain.

J Orthop Res, Vol. 4, No. 4 , 1986


lo-

Statistical Analyses

To evaluate the differences in the mean values of each category, analysis of variance was used to compare data from three or more groups. The Stu- dent's t test was used to compare data from two groups when the population variances of each group could be safely assumed to be equal. When the variances of the two populations were different, we used the method of Welch (48). The F test was used for the evaluation of the equality of the population variances. For comparing the differences in the tensile moduli of the specimens in the two bathing solutions, we used the correlated t test. Most of the statistical and linear regression analyses used in this study were performed with MIDAS, a statistical software package.

H20 A 0.15M NaCl

H2O A 0.15M NaCl

Nr8***

T

RESULTS

Zonal Differences of the Tensile Modulus

In general, from our previous studies, we found that the tensile modulus did not vary in a statistically significant manner from strip to strip; variation from slice to slice was significant. Random variations occurred from joint to joint for each category. The representation of our data reflects this general trend.

Figure 6 shows a typical variation of the intrinsic tensile modulus of the ECM for the four different zones (surface, subsurface, middle, and middle- deep) of the LPGIHWA cartilage samples. The tensile moduli of all the human knee joint cartilage tested were less than 30 MPa; most fell in the range between 1.0 MPa and 15 MPa. For these adult human cartilage, the tensile modulus generally de- creased with increasing distance from the surface. The tensile moduli of the surface zone samples from fibrillated cartilage were significantly diminished as compared with normal cartilage; for LPGI HWA regions p < 0.05, and for LPGiLWA p < 0.01. However, the moduli of the lower zones were very similar to those of normal cartilage. In some fibrillated specimens, particularly with the LPGI LWA specimens, the subsurface zone had significantly greater tensile moduli than those of the surface zone (p < 0.01) (Table 2). For the osteoarthritic specimens, only LWA tissue existed, and the tensile moduli for these specimens were very low, less than 3.0 MPa, and no zonal differences were de- tected (Table 2 ) . In addition, Fig. 6 shows that the

f I k S SS M MD S SS M MD

FIG. 6. Zonal differences of the mean value of the intrinsic tensile moduli of normal and fibrillated adult human knee joint cartilage from the high weight-bearing areas (HWA) of the lateral patellofernoral groove (LPG). The bars represent one standard deviation. Differences of mean intrinsic tensile moduli between deionized H,O and 0.15 M NaCl solution are also shown. Significance of differences between deionized H,O and 0.15 M NaCl are shown as: *I*, p < 0.001; **, p < 0.01; *, p < 0.05; S, surface zone; SS, subsurface zone; M, middle zone; MD, middle-deep zone.

intrinsic tensile moduli of the ECM are greater in deionized water than in 0.15 M NaCl. These differences were statistically significant for cartilage from every zone for both normal and fibrillated joints.

There were no statistical differences or ten- dencies in the variation of the tensile moduli with age in any of the categories of tissues from normal knee joints. However, as the average ages of the fibrillated knee joints and osteoarthritic (TKR) specimens were successively higher, when all the material was taken together there was a trend for a decrease of the tensile modulus with age.

Topographical Differences (HWA vs. LWA)

Table 2 gives the mean values and standard de- viations of the intrinsic tensile moduli of all the specimens tested. In normal cartilage, the LWA specimens for both MFC and LPG locations had higher tensile moduli than the corresponding HWA specimens. This was true for every zone. There

J Orthop Res, Vol. 4, No. 4 , 1986


was no consistent pattern of differences between LPG and MFC, though the LPG/LWA specimens were twice as stiff as the MFCILWA specimens (p < 0.001). However, for fibrillated joint surfaces, no significant differences of the intrinsic tensile moduli were found for cartilage specimens from the surface zone of the HWA and LWA regions. The difference between the tensile moduli of the lower zones of the HWA and LWA became less well defined.

Biochemical Data and Tensile Modulus

Table 3 summarizes all our statistical correlations between the tensile modulus and our biochemical measurements. For normal knee joint cartilage, strong correlations existed, positive and negative, between the tensile modulus and all biochemical measures except water and hydroxyproline/wet weight. All the correlations disappeared, except the collagen/proteoglycan ratio, for cartilage specimens from fibrillated surfaces and no correlations existed for specimens from the osteoarthritic (TKR) cartilage.

Figure 7 shows the correlation between the collagen/proteoglycan ratio and mean values of the tensile modulus for the different zones of normal cartilage, r = 0.714 and p < 0.0001.4 Note that the collagen/proteoglycan ratio was higher for LWA- specimens than for HWA specimens. As expected, the surface zones also had higher collageniproteoglycan ratios that corresponded to their higher tensile moduli. Interestingly, for normal knee joint car-

TABLE 3 . Summary of statistical significance of correlations between intrinsic tensile modulus and

biochemical parameters

Normal Fibrillated Osteoarthritic

Water content Collagen/proteogl ycan C A Hexosamine/uronic acid A Hydroxyproline/dry wt . A Hydroxyprolineiwet wt. Hexosamine/dry wt. B Hexosamine/wet wt. B Uronic acidldry wt. A Uronic acid/wet wt. B

A, p < 0.05; B, p < 0.01; C, p < 0.001. Collagen as calculated assuming 100 hydroxyproline residue in 100,OOO formula weight. Proteoglycan content was calculated assuming 22% uronic acid in the monomer.

Data points represent mean values of LPG and MFC for each zone, i.e., subsurface, middle, and middle-deep zones.

--. 0 HWA A LWA _I’ ‘,I

0 1 2 3 4 5 6 7 Collagen/Proteoglycan (Ratio1

FIG. 7. Relationship between intrinsic tensile modulus and collageniproteoglycan ratio for normal human knee joint cartilage. The variables are related by the linear regression equation E = 4.31 (R) - 9.80 ( r = 0.714, p < 0.001, n = 25) where R = collagen/proteoglycan ratio. The mean values and ?SD of collagen/proteoglycan ratio are 4.00 2 0.20 (n = 4, HWA surface), 4.85 2 1.00 (n = 3, LWA surface), 3.13 2 0.21 (n = 8, HWA middle), and 3.54 2 0.73 (n = 4, LWA middle).

tilage, the correlation between the tensile modulus and hydroxyproline/dry weight was not as strong ( r = 0.486 and p < 0.05). Negative correlations existed between the intrinsic tensile modulus and hexosamine/wet weight, r = -0.563 and p < 0.01, and uronic acidiwet weight, Y = -0.491 and p < 0.05.

For specimens obtained from fibrillated knee joints, the only correlation between the mean value of the tensile modulus for the different zones of the ECM and all of our biochemical measurements was with the collagen/proteoglycan ratio, r = 0.559 and p < 0.05 (Fig. 8).4 For osteoarthritic specimens from TKR patients, no correlations existed between the tensile modulus and any of our biochemical measures. The values of the tensile moduli were very low and the collagen/proteoglycan ratio tended to also be on the low side (Fig. 8).

DISCUSSION

Our equilibrium or flow-independent tensile experiments on human knee joint cartilage showed that the ECM behaves linearly in tension, at least up to 15% strain, and that these intrinsic tensile moduli of the ECM are generally less than 30 MPa.5

Myers (35) measured the intrinsic tensile modulus of normal bovine patellofemoral groove cartilage to be of the order of 1.0 MPa.

J Orthop Res, Vol, 4, No. 4, 1986


OA o /

a

< 12

p -i 10

,L

O o/

Normol A

Collogen/Profeoglycon ( Ao l i o l

FIG. 8. Relationship between the intrinsic tensile modulus and collagen/proteoglycan ratio for normal, fibrillated, and- osteoarthritic (OA) cartilages. For fibrillated tissues, the variables are related by the regression equation E = 1.18(R) - 0.26 ( r = 0.559, p < 0.05, n = 18) where R = collageniproteoglycan ratio. There is no significant correlation noted for OA tissues. The mean value and ?SD of collagen/proteoglycan (Yo weight) are 4.36 2 0.75 (n = 7, normal surface), 6.66 f 2.66 (n = 3, fibrillated surface), 3.64 2 0.87 (n = 4, OA surface), 3.26 5 0.46 (n = 12, normal middle), 4.06 f 1.04 (n = 10, fibrillated middle), and 1.36 2 0.15 (n = 4, OA middle).

These results were significantly lower than those reported by Kempson (22); for some young adults, he reported tensile moduli as high as 150 MPa.6 There are at least two reasons for this observed difference: A flow-generated stiffening effect exists, associated with the constant strain-rate tensile experiment, which tends to stiffen the specimen in tension (25) and our use of a low range of tensile strain in this study (< 15%) where the equilibrium stress-strain relationship is linear (Fig. 5 ) . This linear tensile equilibrium stress-strain relation of the ECM is consistent with the linear compressive equilibrium stress-strain behavior of the ECM and with the theoretical biphasic formulation for articular cartilage (3 1). In addition, this biphasic stiffening effect is very similar to those found in the constant strain-rate, creep, and stress-relaxation compression experiments (24,32). Furthermore, our results for normal tissues do not show any age dependence. We now believe the age dependence of the tensile modulus observed by Kempson (22) might be attributed to surface fibrillation and osteoarthritic changes rather than to some intrinsic aging process causing the decrease of the tensile

Constant strain-rate tensile experiments yield a tensile modulus for the surface zone of young bovine cartilage of about 30 MPa (40,51).

modulus of the ECM. It is to be noted that the “fibrillated specimens” tested were removed from visually normal areas immediately adjacent to areas of overt fibrillation. No evidence of tissue damage exists on these “fibrillated specimens.” Additional studies need to be pursued to ascertain the reasons for the discrepancy between our “age-independent” tensile modulus result for our “normal” cartilage population and Kempson’s “age-dependent’’ tensile modulus result for his “normal” cartilage population (22) .

Normal adult articular cartilage, human and bovine, have very stiff surfaces. This stiffness is prob- ably required to carry the large tensile stress acting parallel (hoop stress) to the articulating surface (30). Stress analysis applicable to the equilibrium condition (elastic model) shows that the loss of this stiff protective superficial layer on fibrillated surfaces will subject the remaining zones of the ECM to unusually high stresses (4), thus causing further damage to the tissue in these lower zones. It is worth noting that in mildly fibrillated surfaces, the tensile moduli of tissues from the lower zones were not diminished; this suggests a progressive degen- erative process starting from the articulating surface. Apparently, osteoarthritic (TKR) cartilage is totally disorganized; the tensile moduli of their ECM are very low and no zonal variations exist.

There are two mechanisms to explain the decrease of the tensile modulus of the ECM with increasing Na+ concentration as measured under the isometric test condition. First, the decrease of the Donnan osmotic swelling pressure associated with an increase of the Na+ in the interstitium will result in an internal release of collagen fibril tension (28). If, as in the isometric experiment, the collagen network is constrained at constant length, this internal release is externally manifested in the decrease of the tensile stress required to maintain the isometric condition used to stretch the specimen (1S,37). Second, an increase of the Na+ in the interstitiurn can cause the ECM to weaken in tension (37) by such mechanisms as charge shielding of the electrically interacting groups between the collagen fibrils and proteoglycans (38,42). Because of this, we believe that the correlation between the intrinsic tensile modulus of the ECM with the collagen/proteoglycan ratio (Fig. 7) is particularly meaningful.

The intrinsic tensile modulus of the ECM for the LWA and HWA regions of the joint confirms another structure-function hypothesis about the function of collagen fibrils and proteoglycans in artic-

J Orthop Res, Vol. 4, N o . 4, 1986


ular cartilage: proteoglycan is the main structural component required to bear the compressive loads for the tissue. Consequently, in lightly loaded regions there should be fewer proteoglycans, thus a higher collageniproteoglycan ratio. This was confirmed biochemically and correlated biomechani- cally (Fig. 7). In addition, for layered structures such as the cartilage/bone system at the joint surface, large tensile hoop stresses are most likely to exist at the periphery of the highly loaded regions. Thus, a higher tensile stiffness is required at the LWA regions.

The statistical correlations between our biochemical and biomechanical data show a most inter- esting result: the intrinsic tensile modulus of the ECM correlated better with collagen/proteoglycan ratio than with hydroxyproline or collagen content. This says that while collagen does contribute to the stiffness of the ECM, it is not the main parameter in determining the stiffness. Rather, the tensile stiffness of the ECM is more strongly affected by the collagen/proteoglycan ratio. This implies that the interactions responsible for the cohesiveness of the solid matrix also contribute significantly to the tensile stiffness of the tissue. Disruption of these interactions, by such mechanisms as increasing the counter-ion concentration in the interstitium of the ECM, will cause a decline of the intrinsic modulus. This has also been confirmed.

The tensile specimens from the fibrillated joint surfaces were taken from areas adjacent to the areas stained with India ink. Visually, the areas from which we took our tensile specimen were normal, i.e., no India ink staining occurred. Yet, our results show that a biomechanical change has occurred; this is evidenced by the pronounced decrease of the intrinsic tensile moduli of specimens harvested from the various surface zones. In addition, there appears to be a stiffening of the immediate subjacent zone associated with this decrease. This might be a biological response of the tissue to the disruptions occurring in the surface zones above by providing another layer of stiff tissue to withstand the loads of articulation. While the collagen/proteoglycan contents for fibrillated cartilage remain comparable to normal cartilage (Fig. 8), the intrinsic tensile moduli of fibrillated cartilage are definitely lower than those of normal cartilage. This and the lack of correlation of the tensile moduli of fibrillated cartilage with any of the other biochemical composition measurements indicate that micro- scopic structural disorganization may be respon-

sible for the inferior tensile properties of the surface zone of fibrillated cartilage.

CONCLUSIONS

1 . The equilibrium tensile stress-strain behavior of the extracellular matrix is linear up to - 15% strain. This flow-independent intrinsic tensile modulus has a value ranging from 1 .O MPa to 30 MPa.

2. The intrinsic tensile moduli of LWA cartilage are always greater than those of HWA cartilage. This corresponds very well with the higher collagen/proteoglycan ratio for LWA cartilage than for HWA cartilage. No consistent pattern of differences exists for the tensile moduli of cartilage from the lateral patellofemoral groove and those from the medial femoral condyle.

3. The strongest positive correlation exists between the intrinsic tensile modulus of the extracellular matrix and the collagen/proteoglycan ratio, followed by a good positive correlation with hydroxyproline/dry weight and a good negative correlation with hexosamine/wet weight.

4. Visually normal regions, not stained by India ink, from fibrillated knee joint surfaces have much lower tensile moduli. At times, the subsurface zones adjacent to the fibrillated regions have higher tensile moduli than the visually normal surface zone.

5 . The interactions between the collagen fibrils and proteoglycans appear to be an impor- tant contributing factor in the tensile stiffness of the extracellular matrix. Disruption of these interactions by charge shielding of possible electro- static interaction sites with increased Na+ concentration within the intersti t ium or osteoarthritic processes weakens the extracellular matrix in tension.

6. The intrinsic tensile modulus of normal articular cartilage does not appear to change with age. Age-related changes in the tensile stiffness appear to be the result of a fibrillation and/or OA process rather than some intrinsic aging process causing the observed decrease of the tensile modulus.

Acknowledgment: This research was sponsored by the National Institute of Arthritis, Diabetes, and Digestive and Kidney Diseases grant AM19094, the Clark and Crossan Endowment at Rensselaer Polytechnic Institute, and the Veterans Administration (Miami), University of Miami. We would like to thank Ms. Rose Boshoff and

J Orthop Re&, Vol. 4 , No. 4 , 1986


Ms. Donna Hutchinson for their help in preparing this manuscript, and Mr. John M. Schoonbeck for his tech- nical assistance. The authors with to express special thanks to Dr. Theodore I. Malinin for providing the human knee joints, without which this study would not have been possible.

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