Abstract
This study describes the first application of a varus loading device (VLD) to the rat hind limb to study the role of sustained altered compressive loading and its relationship to the initiation of degenerative changes to the tibio-femoral joint. The VLD applies decreased compressive load to the lateral compartment and increased compressive load to the medial compartment of the tibio-femoral joint in a controlled manner. Mature rats were randomized into one of three groups: unoperated control, 0% (sham), or 80% body weight (BW). Devices were attached to an animal's leg to deliver altered loads of 0% and 80% BW to the experimental knee for 12 weeks. Compartment-specific material properties of the tibial cartilage and subchondral bone were determined using indentation tests. Articular cartilage, calcified cartilage, and subchondral bone thicknesses, articular cartilage cellularity, and degeneration score were determined histologically. Joint tissues were sensitive to 12 weeks of decreased compressive loading in the lateral compartment with articular cartilage thickness decreased in the peripheral region, subchondral bone thickness increased, and cellularity of the midline region decreased in the 80% BW group as compared to the 0% BW group. The medial compartment revealed trends for diminished cellularity and aggregate modulus with increased loading. The rat-VLD model provides a new system to evaluate altered quantified levels of chronic in vivo loading without disruption of the joint capsule while maintaining full use of the knee. These results reveal a greater sensitivity of tissue parameters to decreased loading versus increased loading of 80% BW for 12 weeks in the rat. This model will allow future mechanistic studies that focus on the initiation and progression of degenerative changes with increased exposure in both magnitude and time to altered compressive loads.
Introduction
Clinically, malalignment of the knee and increased body mass index are associated with development of primary osteoarthritis (OA) [1,2]. Increased articular contact stresses are predictive of OA development [3], while procedures aiming to redistribute or reduce articular stress have been utilized to alleviate pain associated with OA [4], including osteotomy, joint distraction, bracing, and corrective footwear. Requisite loading maintains articular cartilage homeostasis, while abnormal levels of loading result in cell death and matrix degradation [5,6]; however, the thresholds of noninjurious sustained loading have yet to be defined. Commonly used small animal models of OA alter joint mechanics via transection of ligaments and/or menisci [7]. These approaches compromise the joint capsule, typically alter the cartilage contact mechanics in an undefined manner, and may be more reflective of the development of posttraumatic OA rather than the slower development of primary OA.
Recently, a varus loading device (VLD) applied to the hind limb of a rabbit has been described to study the effects of controlled levels of altered compressive loading across the tibio-femoral joint in a compartment-specific manner [8–10]. The VLD applies altered loads in addition to the normal loads across the joint without disruption of the joint capsule while maintaining full use and range of motion of the joint. The VLD-animal model allows the role of altered mechanical loading, a main risk factor for the development of primary OA, to be isolated and its contribution to the initiation and progression of degenerative joint changes evaluated in vivo.
This pilot study describes the application of the VLD to the rat to investigate load-induced alterations in response to in vivo chronic varus loading of the rat tibio-femoral joint. We hypothesized that 12 weeks of increased loading of the medial compartment and decreased loading of the lateral compartment would induce compartment-specific changes to the material properties (cartilage aggregate modulus, permeability, and subchondral bone modulus) and structure (cartilage cellularity, degeneration score, and tissue thickness) of the tibial plateau.
Methods
Animal Model
Thirteen, 9-month-old, male, Sprague-Dawley rats (693 ± 32 g) were randomly assigned to control (n = 5), sham: 0% body weight (BW) (n = 4), or 80% BW groups (n = 4). Rats were housed in single cages (19(w) × 32(l) × 19(h) cm), fed chow (Prolab RMH-3000, Purina, St. Louis, MO, USA), and maintained on a 12:12 h light/dark cycle. All procedures were carried out in accordance with the Institutional Animal Care and Use Committee. Rats in the 0% and 80% BW groups underwent surgery to attach custom transcutaneous bone plates to the lateral aspect of the left tibia and femur, as previously described [11].
Following a 2-week recovery, these rats were fit with a VLD (Fig. 1), the spring torque set to apply the target change in contact load to the tibial plateau, and the VLD engaged. Engagement of the VLD produced a varus moment about the tibio-femoral joint, which increased the compressive force in the medial compartment and decreased the compressive force in the lateral compartment by an equivalent amount. As determined from free body analysis, the change in contact load (ΔP) is ΔP = Fs* L2/d, where Fs = medial directed force at distal tibia, d = intercompartmental moment arm, and L2 = tibia moment arm (Fig. 1(c)). The total weight of the attached bone plates and VLD was less than 2% BW. Target loads (0% and 80% BW) were applied 12 h/day during the 12-h dark cycle, when rat activity is the highest, 7 days/week for 12 weeks. The load applied to the distal tibia was measured twice daily using a handheld load cell (Shimpo FGE-0.5x, Cedarhurst, NY, USA) and adjusted if beyond ±10% of the target level. Following the 12-week loading protocol, animals were euthanized and tibial osteochondral specimens harvested from experimental and contralateral limbs and stored at –80° C.
Fig. 1.
Application of the varus loading device (VLD) to the rat. (a) Lateral view of the VLD applied to rat femur and tibia via transcutaneous bone plates secured with bone screws. The axis of the bearing is aligned with the epicondylar axis of the femur utilizing fluoroscopy. Setting the torque of the torsion spring applies a varus moment to the distal tibia, which results in an increased compressive load (+ΔP) in the medial compartment (M) and decreased load (–ΔP) in the lateral compartment (L) of the tibio-femoral joint. (b) Lateral view of the VLD disengaged. To disengage the VLD, the offset link and a portion of the load link are removed and the remaining load link is rotated into alignment with the femur tube and secured, thus removing the varus moment. (c) Anterior-posterior view of rat femur and tibia with a VLD attached and engaged, illustrating the delivery of a compressive overload to the medial (M) compartment and a decrease in loading in the lateral (L) compartment. ((d) and (e)) Rat with VLD attached and engaged. (Adapted from Roemhildt et al. [11]).
Mechanical Evaluation
The material properties (aggregate modulus, permeability, and Poisson's ratio) of the articular cartilage were evaluated using a biphasic creep-indentation test [12,13]. Central sites in the medial and lateral compartments of the tibial plateau (Fig. 2(a)) were tested as previously described with a custom materials testing device utilizing a cylindrical, plane-ended, porous, 0.5-mm-diameter indenter tip and phosphate buffered saline bath with protease inhibitors [14]. Following application of a tare load (0.044 MPa) for 15 min, the indentation test proceeded with the application of the test load (0.125 MPa) with data sampled at 1 Hz until displacement reached equilibrium (less than a 0.5 -μm change in displacement over a 300-s period) (mean strain = 3.9 ± 1.6%). Material properties of the articular cartilage were determined by curve-fitting the load-displacement response with the biphasic indentation creep solution via a nonlinear regression procedure [13,14].
Fig. 2.
(a) Superior view of left tibia plateau illustrating locations of mechanical test sites (•) in the medial and lateral compartments of the tibial plateau. (b) Safranin-O- and Fast Green-stained section illustrating identification of the articular cartilage surface (black), tidemark (dark green), calcified cartilage/subchondral bone boundary (light green), and subchondral bone plate (red) at 61 points (gray vertical lines) across each compartment. Yellow vertical lines divide the compartment into three equal-width regions (peripheral, central, and midline).
The local compressive modulus of the subchondral bone underlying the cartilage test site was determined using a microindentation test [15]. Three cycles of repeated loading were applied (2.452 N, 30 s hold) to the osteochondral specimen via a stainless steel needle with a parabolic-shaped tip. The loads on and displacements of the indenter were measured at 128 Hz, and the slope of the load-displacement response during the unloading portion of the third cycle was used to calculate the subchondral bone modulus.
Histological Analyses
Specimens were formalin-fixed, decalcified with 10% ethylenediamine tetra-acetic acid, and paraffin-embedded [16]. Serial 5 -μm coronal sections were prepared from the posterior half of the tibial plateau at 200 -μm intervals [7] and stained with Safranin-O and Fast Green or hematoxylin and eosin (H&E) [17] prior to examination under a light microscope (BX50, Olympus Inc.) fit with a digital camera (RET-2000R-F-CLR-12-C, QImaging, Surrey, BC, Canada) for acquisition of images (1200 × 1600 pixels). Three consecutive sections were evaluated for each plateau using custom Matlab (MathWorks, Natick, MA, USA) code to facilitate measurements. The articular surface, tidemark, calcified cartilage-subchondral bone junction, and inferior boundary of the subchondral bone plate were identified on Safranin-O slides to determine articular cartilage, calcified cartilage and subchondral bone thicknesses, and articular cartilage area (Fig. 2(b)). Region-specific measures were calculated by dividing each compartment into 3 regions of equal width (peripheral, central, and midline; Fig. 2(b)). Cellularity was calculated as the number of chondrocytes per articular cartilage area, and all chondrocytes with visible nuclei were counted on a composite of digital images of H&E stained sections collected using a 10X objective. Degeneration score was evaluated utilizing the Osteoarthritis Research Society International (OARSI) recommendations for the rat [7]. The cartilage matrix loss width (at the surface, midzone, and tidemark), cartilage degeneration, total cartilage degeneration width, significant cartilage degeneration width, zonal-depth ratio of lesions, and osteophyte size were determined as detailed in the Appendix (Table 2).
Table 2.
OARSI cartilage degeneration score for the rat (adapted from Gerwin et al. [7])
| Outcome measure | Description | |
| 1 | Cartilage matrix loss width | The width of areas of complete collagen matrix loss measured along the surface (0%), midzone (50%), and tidemark. Widths were normalized by the width of the entire compartment. |
| 1a. Width matrix loss surface (%) | ||
| 1b. Width matrix loss midzone (%) | ||
| 1c. Width matrix loss tidemark (%) | ||
| 2 | Cartilage degeneration score | Collagen matrix fibrillation/loss and chondrocyte death/loss are the principal determinants of the score. The compartment was divided into three regions of equal width: periphery (P), central (C), and midline (M). The original surface of the tissue was estimated, and the percent area of each region exhibiting loss of chondrocytes (>50% loss of normal cellularity) or loss of matrix was measured. A score was assigned to each region as follows: |
| 2a. Cart degeneration_peripheral (%) | 0: No degeneration | |
| 2b. Cart degeneration_central (%) | 1: Minimal degeneration; 5%–10% of the total projected cartilage area affected by matrix or chondrocyte loss | |
| 2c. Cart degeneration_midline (%) | 2: Mild degeneration; 11%–25% affected | |
| 2d. Total cart degeneration (%) | 3: Moderate degeneration; 26%–50% affected | |
| 2e. Total cart degeneration (score) | 4: Marked degeneration; 51%–75% affected | |
| 5: Severe degeneration; greater than 75% affected | ||
| The cartilage degeneration score was calculated by summing values obtained for each zone. Maximum cartilage degeneration score = 15. | ||
| 3 | Total cartilage degeneration width | The total width of the area of articular cartilage affected by any type of degenerative change (matrix fibrillation/loss, proteoglycan loss with or without chondrocyte death) is measured at the surface. This measurement takes into account foci of minor change (proteoglycan loss only). The widths of all degenerated areas are summed and normalized by the width of the compartment. |
| 3. Width cart degeneration (%) | ||
| 4 | Significant cartilage degeneration width | Measurement of the width of the tibial cartilage in which 50% or greater of the thickness (from surface to tidemark) is seriously compromised by chondrocyte loss (>50%) with or without matrix loss |
| 4. Width sig. cart degeneration (%) | ||
| 5 | Zonal (regional) depth ratio of lesions | Measurement of the depth of cartilage degeneration as indicated by chondrocyte and proteoglycan loss (independent of matrix fibrillation) taken at the midpoint in each of the three regions across the compartment. The depth ratio is calculated by dividing the depth of the area of degeneration by the thickness of the cartilage (from projected cartilage surface to tidemark). |
| Supplemental measure | ||
| 5a. Depth ratio_peripheral | ||
| 5b. Depth ratio_central | ||
| 5c. Depth ratio_midline | ||
| 6 | Osteophytes | Measurement of thickness from chondro-osseous junction to overlying cartilage surface at thickest point and converted to grade: |
| 0: <200 μm; 1: small 200–299 μm; 2: moderate 300–399 μm; 3: large 400–499 μm, 4: very large ≥ 500 μm |
Statistical Analyses
Analyses of variance were used to evaluate differences in outcome measures across treatment conditions (control, 0%, 80% BW). For histological measures, the aggregate mean of three consecutive slides were used for analyses. Separate analyses were performed for each compartment (medial and lateral) and leg (experimental and contralateral). Following a significant F-test [18], pairwise comparisons using Fisher's least significant difference procedure were used to test for differences between experimental groups. Statistical analyses were performed using sas statistical software version 9 (sas Institute, Cary, NC, USA).
Results
Gross observation of the tibial plateau revealed minimal erosion or fibrillation of the articular cartilage in all experimental groups. Load-induced alterations to the tibial plateau were most pronounced for region-specific thickness measures and cellularity in the lateral compartment of the experimental limb, which experienced decreased loading (Table 1). The thickness of the articular cartilage in the peripheral region of the lateral compartment decreased 19% in the 80% BW group as compared to the 0% BW group (p = 0.04) and 26% as compared to the control group (p = 0.01). In contrast, the thickness of the subchondral bone increased 38% in the 80% BW group as compared to the 0% BW group (p = 0.01). A 23% reduction in the cellularity of the midline region of the lateral compartment was observed in the 80% BW group as compared to the 0% BW group (p = 0.04). No significant differences across groups were observed for material properties or components of the OARSI score for the lateral compartment. Differences between control and 0% BW groups were observed for articular cartilage thickness (central region), subchondral bone thickness, and cellularity (midline region), as detailed in Table 1.
Table 1.
Material properties and histological outcome measures for the medial and lateral compartments of the tibial plateau of the experimental limb
|
Lateral compartment |
Medial compartment |
|||||||
|---|---|---|---|---|---|---|---|---|
| Outcome measure | Control | 0% BW | 80% BW | ES | Control | 0% BW | 80% BW | ES |
| Thickness | ||||||||
| Cartilage thickness (μm) | 239 (30) | 214 (14) | 204 (36) | –0.37 | 302 (30) | 267 (6) | 293 (21) | 1.68 |
| Peripheral (μm) | 163 (17) | 149 (15) | 121 (19)a , b | –1.64 | 198 (24) | 179 (16) | 182 (28) | 0.13 |
| Central (μm) | 306 (12) | 262 (28)c | 256 (44)b | –0.16 | 325 (29) | 302 (13) | 322 (34) | 0.78 |
| Midline (μm) | 254 (54) | 234 (27) | 240 (49) | 0.15 | 383 (60) | 325 (30) | 376 (55) | 1.15 |
| Calcified cartilage (μm) | 112 (21) | 107 (14) | 118 (20) | 0.64 | 69 (11) | 72 (9) | 67 (18) | –0.35 |
| Subchondral bone (μm) | 161 (26) | 127 (14)c | 175 (21)a | 2.51 | 271 (87) | 250 (70) | 222 (47) | –0.47 |
| Cellularity | ||||||||
| Cellularity overall mean (cells/mm2) | 592 (78) | 771 (96) | 659 (138) | –1.15 | 406 (97) | 426 (68) | 319 (93) | –1.31 |
| Cellularity peripheral (cells/mm2) | 724 (131) | 846(154) | 846 (171) | 0.00 | 649 (114) | 652 (107) | 533 (110) | –1.10 |
| Cellularity central (cells/mm2) | 514 (88) | 708 (146) | 600 (138) | –0.76 | 323 (86) | 344 (69) | 222 (101) | –1.41 |
| Cellularity midline (cells/mm2) | 609 (65) | 815 (126)c | 628 (137)a | –1.42 | 358 (144) | 387 (80) | 301 (108) | –0.90 |
| OARSI scoring | ||||||||
| 1a. Width matrix loss surface (%) | 41.1 (13.6) | 38.1 (24.6) | 17.2 (12.2) | –1.08 | 15.7 (11.2) | 29.6 (12.1) | 34.2 (25.6) | 0.23 |
| 1b. Width matrix loss midzone (%) | 1.8 (3.7) | 0.4 (0.7) | 0.0 (0.0) | –0.81 | 0.0 (0.0) | 0.16 (0.33) | 0.19 (0.38) | 0.08 |
| 1c. Width matrix loss tidemark (%) | 0.0 (0.0) | 0.0 (0.0) | 0.0 (0.0) | 0.0 (0.0) | 0.44 (0.88) | 0.0 (0.0) | –0.71 | |
| 2a. Cart degeneration peripheral (%) | 8.08 (8.6) | 2.3 (2.1) | 3.0 (2.7) | 0.29 | 21.4 (17.6) | 23.0 (2.9) | 35.5 (18.4) | 0.95 |
| 2b. Cart degeneration central (%) | 25.9 (14.0) | 16.8 (9.8) | 26.3 (10.5) | 0.94 | 46.3 (36.3) | 64.7 (9.0) | 80.7 (24.7) | 0.86 |
| 2c. Cart degeneration midline (%) | 22.3 (10.0) | 14.5 (4.9) | 18.3 (9.6) | 0.50 | 46.4 (30.2) | 51.6 (20.5) | 64.0 (20.4) | 0.61 |
| 2d. Total cart degeneration (%) | 21.3 (10.1) | 13.2 (5.6) | 19.1 (8.2) | 0.84 | 41.9 (28.9) | 51.5 (11.9) | 67.4 (17.0) | 1.08 |
| 2e. Total cart degeneration (score) | 5.1 (1.8) | 3.8 (1.6) | 4.3 (1.8) | 0.29 | 7.9 (4.0) | 9.8 (1.3) | 11.4 (2.1) | 0.92 |
| 3. Width cart degeneration (%) | 65.3 (9.0) | 62.4 | 67.3 (6.8) | 0.27 | 42.2 (14.3) | 49.4 (19.4) | 51.8 (23.4) | 0.11 |
| 4. Width sig. cart degeneration (%) | 3.7 (5.3) | 0.0 (0.0) | 1.0 (1.5) | 0.94 | 18.0 (24.3) | 16.0 (12.5) | 35.3 (15.9) | 1.35 |
| 5a. Depth ratio peripheral | 0.09 (0.07) | 0.13 (0.22) | 0.17 (0.10) | 0.23 | 0.05 (0.04) | 0.15 (0.26) | 0.07 (0.05) | –0.43 |
| 5b. Depth ratio central | 0.10 (0.04) | 0.10 (0.09) | 0.14 (0.07) | 0.50 | 0.03 (0.02) | 0.05 (0.15) | 0.08 (0.08) | 0.25 |
| 5c. Depth ratio midline | 0.16 (0.08) | 0.19 (0.12) | 0.22 (0.09) | 0.28 | 0.08 (0.04) | 0.21 (0.15) | 0.12 (0.10) | –0.71 |
| 6. Osteophyte score | 0.0 (0.0) | 0.0 (0.0) | 0.0 (0.0) | 0.0 (0.0) | 0.0 (0.0) | 0.0 (0.0) | ||
| Material properties | ||||||||
| Aggregate modulus (MPa) | 1.63 (0.52) | 1.96 (0.37) | 1.56 (0.29) | –1.20 | 2.05 (0.80) | 2.03 (0.50) | 1.41 (0.24) | –1.58 |
| Permeability (10–16 m4/Ns) | 7.8 (9.9) | 2.1(1.2) | 3.4 (1.9) | 0.82 | 3.28 (1.85) | 2.67 (2.48) | 4.18 (2.07) | 0.66 |
| Poisson's ratio | 0.28 (0.13) | 0.27 (0.11) | 0.25 (0.14) | –0.16 | 0.29 (0.09) | 0.33 (0.08) | 0.27 (0.10) | –0.66 |
| Bone modulus (GPa) | 3.01 (1.29) | 3.44 (1.35) | 3.03 (0.61) | –0.39 | 3.35 (0.55) | 3.45 (0.70) | 4.38 (2.89) | 0.44 |
Note: Data presented as mean (standard deviation).
ES = Cohen's d effect size between 0% BW and 80% BW groups.
Significant difference between 0% BW vs. 80% BW groups (p < 0.05).
Significant difference between control vs. 80% BW groups (p < 0.05).
Significant difference between control vs. 0% BW groups (p < 0.05).
In the medial compartment of the experimental limb, no overall significant differences across groups were detected. Pairwise comparisons revealed a trend for diminished cellularity (35%) in the 80% BW group as compared to the 0% BW group in the central region (p = 0.09; Fig. 3). Similarly, the aggregate modulus of the articular cartilage of the 80% BW group was 30% less than that of the 0% BW group (p = 0.09).
Fig. 3.
Representative H&E stained sections from 0% BW (a) and 100% BW (b) groups, suggesting diminished chondrocytes with increased loading of the medial compartment
No significant differences across groups were observed for outcome measures in the contralateral limb.
Discussion
This work demonstrates the feasibility of applying the VLD to the rat and delivering controlled chronic-load alteration of 80% BW to the tibio-femoral joint for 12 weeks without disrupting the joint capsule. The lateral compartment experienced diminished compressive loading, which resulted in decreased cartilage thickness, increased subchondral bone thickness, and decreased cellularity that preceded significant alteration of material properties and cartilage degeneration, whereas the medial compartment experienced increased loading without producing significant differences across groups over 12 weeks. These results indicate that articular cartilage and subchondral bone may be more sensitive to decreased rather than increased chronic compressive loads over 12 weeks in the rat.
In vivo studies evaluating the effects of diminished loading are limited with disparate findings. Our present findings of decreased articular cartilage thickness and cellularity and increased subchondral bone thickness with diminished loading of the lateral compartment are consistent with results from a rat model, where joint unloading was achieved by tail suspension with articular cartilage thickness decreased and calcified cartilage increased following one month of treatment [19]. In contrast, previous work utilizing the rabbit VLD model found articular cartilage thickness increased following 24 weeks of decreased loading, with no significant changes in the thickness of the subchondral bone [10]. Diminished loading via osteotomy in the guinea pig was associated with increased proteoglycan and collagen content [20]. Clinically, articular cartilage thinning was observed in knee cartilage following reduced joint loading subsequent to spinal cord injury [21], whereas increased tibio-femoral joint space width was observed at one year follow up following two months of unloading via joint distraction [22]. The effects of diminished loading on the morphology of joint tissues suggest that effects may be species, load magnitude, and duration dependent.
In previous studies utilizing a rabbit VLD model with levels of altered loading ranging from 22%–80% BW over 12–24 weeks, load-induced changes were more prominent in the medial tibio-femoral compartment, which experienced an increase in joint loading versus the lateral compartment [8,10]. This may be due to differences in outcome measures (mechanical versus histological); sensitivity of species-specific techniques to determine outcome measures; relative age; duration and magnitudes of applied loading; and species-related differences in tissue properties, activity level, anatomy, joint loading, and gait. There is consistency across all our work in that, at 12 weeks, no significant differences in material properties of the articular cartilage in the lateral compartment were observed. Furthermore, the observed trend for diminished cellularity with increased load in the medial compartment of the rat is similar to load-induced chondrocyte loss observed in prior rabbit studies [8,10].
The chronic load alteration applied to the tibio-femoral joint was in addition to the normal dynamic loads produced across the knee by muscle contractions and body weight during activities with a magnitude, in relation to body weight, approximately twice that experienced in the human knee with varus alignment [23]. Altered chronic loading of 80% BW may be relatively small in comparison to the magnitude of peak dynamic loads produced during ambulation. Although a reduction in vertical ground reaction force was observed with attachment of the VLD to the rat, no additional differences were observed when altered loads were applied [11]. Since we have previously observed a sham effect with surgery for VLD attachment [8,11], we focused on comparison of the 0% BW group to the 80% BW group to isolate the effects of altered loading.
While statistical comparisons of outcome measures between medial and lateral compartments were not performed in this study, our data suggest that compartment-specific differences may exist for select outcome measures, as we have previously reported for the material properties of rabbit articular cartilage [14] and the cellularity of rat articular cartilage [24].
The response of the articular cartilage to altered loading of 80% BW is less severe than changes observed in commonly used transaction-based models of posttraumatic OA in the rat, in which erosion develops within 3–12 weeks [7]. An animal model in which degenerative changes develop gradually in response to controlled loading may be more relevant to primary OA in humans. The VLD model may be particularly useful to study the contribution of altered loading to the early events occurring with joint degeneration, such as chondrocyte loss, which occurs early in the human OA disease process [25,26].
Though few outcome measures reached statistical significance in this pilot study, large effect sizes (>0.8) were observed with loading for key outcome measures, including cellularity and aggregate modulus (Table 1). This work demonstrates feasibility of the rat VLD model to study the effects of in vivo chronic load alteration, which can be utilized in future studies to determine if gross degenerative changes of the joint structures develop with increased exposure to or increased magnitudes of altered loading or if the observed changes may be ameliorated if the altered loading is removed.
The rat VLD model provides a system to evaluate the application of quantified levels of chronic, compressive load alteration to the tibio-femoral joint in vivo. Preliminary results indicate the sensitivity of joint tissues in the lateral compartment to diminished loading over the relatively short time interval of 12 weeks. The rat VLD model will allow future evaluation of the effects of increased exposure to quantifiable altered loads of increasing magnitude and duration and the mechanisms by which degenerative changes initiate.
Acknowledgment
The authors are grateful to Calsey Grant, Dr. Nelson Tacy, Amy Gassman, and Danielle Funaro for their assistance with experimental procedures and to Synthes for providing the bone screws used in this work. Funded by NIH AR052815. The work performed was not influenced by the funding provided. The authors of this work have no conflicts of interest.
Appendix
Contributor Information
M. L. Roemhildt, e-mail: maria.roemhildt@med.uvm.edu.
K. Anderson, McClure Musculoskeletal Research Center, , Department of Orthopedics and Rehabilitation, , College of Medicine, , University of Vermont, Burlington, VT, 05405
G. J. Badger, Department of Medical Biostatistics, , University of Vermont, Burlington, VT, 05405
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