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. Author manuscript; available in PMC: 2009 Dec 1.
Published in final edited form as: J Biomed Mater Res A. 2009 Dec;91(3):795–805. doi: 10.1002/jbm.a.32251

A Novel Biomimetic Polymer Scaffold Design Enhances Bone Ingrowth

Chris P Geffre 1, David S Margolis 1, John T Ruth 1, Donald W DeYoung 2, Brandi C Tellis 1, John A Szivek 1
PMCID: PMC2767470  NIHMSID: NIHMS113402  PMID: 19051300

Abstract

There has been recent interest in treating large bone defects with polymer scaffolds because current modalities such as autographs and allographs have limitations. Additionally, polymer scaffolds are utilized in tissue engineering applications to implant and anchor tissues in place, promoting integration with surrounding native tissue. In both applications, rapid and increased bone growth is crucial to the success of the implant. Recent studies have shown that mimicking native bone tissue morphology leads to increased osteoblastic phenotype and more rapid mineralization. The purpose of this study was to compare bone ingrowth into polymer scaffolds created with a biomimetic porous architecture to those with a simple porous design. The biomimetic architecture was designed from the inverse structure of native trabecular bone and manufactured using solid free form fabrication. Histology and μCT analysis demonstrated a 500-600% increase in bone growth into and adjacent to the biomimetic scaffold at five months post-op. This is in agreement with previous studies in which biomimetic approaches accelerated bone formation. It also supports the applicability of polymer scaffolds for the treatment of large tissue defects when implanting tissue-engineering constructs.

Keywords: scaffolds, μCT, histomorphometry, biomimetic, polybutylene terephthalate

Introduction

Bone allografts and autografts are the current standard of care for treating bone defects1-5. Both treatments have drawbacks since they rely on limited supplies of material, donor site morbidity6-10 and the possibility of disease transmission or immune rejection with allografts11,12. Metal implants have also been used to treat bone defects13. While metal implants are readily available and eliminate the risk of transmittable diseases from donor to host, they can necessitate additional surgeries due to complications such as stress shielding, implant failure or infection14-16. There has been recent interest in resorbable polymer scaffolds which encourage bone regeneration to treat bone defects17. These scaffolds can be prepared prior to or during surgery, can be modified to alter mechanical strength and resorption rate, can be created in custom shapes unique to each defect site, and can be produced with highly controlled structures.

Polybutylene terephthalate (PBT) polymer based scaffolds have been used previously to treat bone defects, particularly as a copolymer blended with polyethylene oxide (PEO) where it has been noted to encourage bone growth in animal models18-25. Previous studies have shown that altering the ratio of PBT to PEO in these scaffolds changes its osteoconductive properties and resorption rate20,21,23,24. While extensive animal work has indicated accelerated bone defect repair with PBT/PEO copolymers20,21,23, a recent clinical study has shown that the copolymer is not osteoconductive in patients26. This suggests that alternative approaches to inducing bone formation in PBT based implants are necessary. Pure PBT scaffolds with osteoconductive calcium phosphate ceramic (CPC) particles and/or osteoinductive protein coatings have also been used in animal models to support bone growth27-29 and may be an alternative to using PBT copolymers for the repair of large bone defects. Additionally, CPC particles have been shown in previous studies to be osteoconductive in patients30. Pure PBT scaffolds are stiffer than PBT/PEO copolymers and have a slower resorption rate24, allowing PBT scaffolds to maintain mechanical integrity for long periods after surgical placement. This may be valuable to allow better integration of tissues on and around the scaffold.

In addition to repairing large bone defects, the slow resorption and long-term mechanical integrity of PBT scaffolds provide a material that can be used to repair tissue interfaces and with the addition of sensors can be used to monitor tissue loading. PBT scaffolds placed in load-bearing locations, such as subchondral bone in the knee, are able to support bone attachment and facilitate limited cartilage healing28, demonstrating that these scaffolds can be used to repair damage at musculoskeletal tissue interfaces. Such scaffolds may be crucial to successfully repairing focal cartilage defects with tissue engineered constructs as the long-term mechanical support provided by PBT will allow sufficient time for engineered tissues to integrate with surrounding native tissues. Furthermore, sensate PBT scaffolds have been developed by attaching strain sensors to their surfaces and have been utilized to collect long-term in vivo joint loads27,28,31. Sensate scaffolds may offer additional research and clinical applications compared to traditional scaffolds as they can be used to monitor healing as well as study, modulate and test the efficacy of rehabilitation. They can also provide a method to detect implant weakening and warn the patient prior to catastrophic failure.

In addition to altering the ratio of copolymers or applying osteoconductive coatings to encourage bone growth, previous studies have shown that pore size, structure and interconnectivity influence bone growth32-37. While many studies have looked at the effect of pore size and connectivity on bone ingrowth32-41, showing that pore sizes ranging from 100 - 400μm support the most extensive bone growth38-41, few studies have investigated the effect of pore shape on bone growth. One study demonstrated that square pores support more bone growth than round pores and attributed this difference to the cross-sectional area of the pores33. PBT scaffolds provide an ideal material to test how different scaffold designs effect bone growth because the implants can be made with solid free form fabrication27-29,31 providing a high degree of control over the exact pore shape, size, and interconnectivity. Recently, CT and MRI data sets have been used in combination with free form fabrication to produce scaffolds that exactly replicate the native environment in which they will be implanted42-44.

The goal of this study was to compare bone ingrowth into scaffolds with a biomimetic inverse trabeculated pore structure, to previously studied simple porous scaffolds27,28. The inverse trabeculated pore structure was created from inverted μCT data, containing solid regions in place of marrow spaces and empty space in place of bone tissue, so that bone growth into the pores would re-create the natural structure of bone once the scaffold has been resorbed. Since creating scaffolds with a biomimetic pore structure unique to each patient is a labor-intensive process and necessitates exposing the patient to radiation during μCT imaging, it is important to determine whether this new approach to scaffold design does accelerate bone growth and justify the additional expense and patient risk.

Methods and Materials

Scaffold design

A simple porous scaffold was designed according to a published procedure45. Briefly, the scaffold was designed in SolidWorks (Concord, MA) with a solid cylindrical outer skin and an interconnected porous interior. A solid horizontal layer of PBT simulating subchondral bone was placed between the cylindrical structure and the domed surface, which interfaced with the adjacent joint surface (Figure 1). The interconnected porous interior was created using a grid pattern which was rotated 45° in each layer in order to create a tortuous porous construct intended to enhance bone ingrowth (Figure 2). The solid outer cylinder was utilized to provide structural strength to the scaffold and it allowed subsequent strain gauge attachment. The outer scaffold skin was designed with three sets of openings (three holes measuring 1mm in diameter) spaced equidistant along the longitudinal axis to allow bone ingrowth into the interior of the scaffold from the lateral edges. The solid layer of PBT was created between the porous interior and domed surface to prevent bone growth into the joint space. The SolidWorks design was imported into a Windows (Microsoft, Redmond, WA) based computer which controlled a FDM 1650 modeler (Statasys, Eden Prairie, MN) and the scaffold was manufactured one horizontal layer at a time from PBT filament that was extruded through a heated tip .3049mm in diameter. The finished scaffold was 8.9mm in diameter and 11.3mm in length.

Figure 1.

Figure 1

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Schematic illustration depicting the sections of the scaffold including the dome, solid outer skin, horizontal layer, and porous interior.

Figure 2.

Figure 2

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CT images representing the changes in inner porous architecture. (A) illustrates the grid like pattern that was rotated 45° between layers while (B) represents the biomimetic inverse trabeculated porous structure.

The inverse trabecular pore structure for the biomimetic scaffold was created by first taking one 17mm diameter bone core from a cadaveric canine femur similar in size to the joints of the test animals selected for the in vivo portion of the experiment. The core was scanned using a μCT scanner (Scanco, Bassersdorf, Switzerland) at 18μm resolution and a region of interest in the trabecular bone was selected. The data set was modified to create an inverse trabecular pattern where marrow spaces were filled in with polymer and trabecular spaces were left as pores to allow bone ingrowth (Figure 2). Similar to the construction of the simple porous scaffold, a cylindrical outer skin was added and a solid layer of PBT was placed between the dome and the cylindrical porous section. Additionally, the dome was designed to support tissue-engineered cartilage. Three equidistantly spaced windows (each measure 2.95 by 6.77mm) were placed in the outer skin to allow for bone ingrowth. This file set was exported as a stereolithography (STL) file to the FDM 1650 modeler, which built the scaffold layer by layer.

Scaffold Characterization

Six scaffolds of each type were scanned in the μCT scanner at 18μm resolution and evaluated for porosity using a threshold of 50. Scaffolds were then axially compressed on a servo-hydraulic materials testing system (MTS Corp, Minneapolis, MI) to determine stiffness of each scaffold type. Scaffolds were loaded at a displacement rate of 4.2mm/min with no preconditioning or preload. For each scaffold, load (kgf) and stroke (mm) were collected using custom data collection software (Labview 5.0.1, National Instruments, Austin, TX) running on a Macintosh G4 computer. Stiffness was calculated from the slope of the each scaffold's linear stress-strain curve.

Scaffold Preparation for In Vivo Placement

Following production of scaffolds to be implanted, three 1000ohm strain gauges (Micro-Measurements, Raleigh, NC) were attached to four scaffolds of each type along the longitudinal axis of each scaffold (Figure 3) in order to measure in vivo loads during gait. Loads have been reported in a separate manuscript27. Four additional scaffolds of each type were left unwired to allow μCT data to be collected and compared to histology measurements. All scaffolds were coated according to a published procedure27,28,46-48 in a mixture of calcium phosphate ceramic particles, previously designated CPC 6 and 730,49. This mixture of CPC's has been shown to encourage rapid bone growth in a canine model as well as in patients30,49. Wired scaffolds were then calibrated using a servo-hydraulic materials testing system to create a unique load/strain calibration curve for each wired scaffold that could be used to interpret in vivo strain measurements as loads27,28,31. Each scaffold was sterilized using ethylene oxide and aerated prior to the surgery. The morning of the surgeries scaffolds were coated with 1μg of TGF-β1 to encourage bone ingrowth50.

Figure 3.

Figure 3

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A strain-gauged scaffold showing one of the strain gauges, which was attached to the scaffold for load monitoring. Also illustrated is the dome of the scaffold along with the open window and biomimetic interior.

Scaffold Placement

Two groups of four hounds were selected to undergo implantation of either the simple porous scaffolds or the biomimetic scaffolds in both femoral condyles of the right stifle. All animals were cared for and operated on according to NIH guidelines (NIH Publication 82-23, Rev. 1985).

An orthopaedic surgeon (J.T.R.) performed all of the scaffold implantation surgeries according to previously published procedures27,28,31. Briefly, an incision was made in the skin and the medial condyle of the right stifle was exposed using blunt dissection. A guide drill was used to create a 3mm hole through the condyle. The guide was drilled through the lateral middiaphysis of the femur to allow wires from the strain gauges to pass out through the lateral aspect of the femur. A 9mm reamer was placed over the guide and used to create a hole in which the scaffold was press fit (Figure 4). The scaffold was orientated in line with the femur's axis of loading so that at 30° of stifle flexion the longitudinal axis of the scaffold was perpendicular with the ground (Figure 5), and was implanted with the domed surface flush with the surrounding cartilage tissue. Scaffolds were implanted into the lateral condyle using the same procedure, but without drilling and passing wires through the mid diaphysis of the femur because they had no sensors and as such were left unwired.

Figure 4.

Figure 4

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Photographs during scaffold implantation depicting the guide drill (A) and the 9mm reamer (B) which was subsequently used to create the hole in which the scaffold was press fit.

Figure 5.

Figure 5

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Schematic illustration showing the location of the scaffolds placed in the medial and lateral condyles.

Animal Care and Sample preparation

Animals were returned to limited load bearing activities by one week post-op and were returned to a full running regimen (at least once per week) after two weeks post-op according to a previous published procedure27. Test animals were labeled with tetracycline for 3 days at 18 days prior to euthanasia and 4 days before euthanasia. They were labeled every 8 hours during the two 3 day sessions, which allowed calculation of bone formation rate. At 5 months post-op the test animals were euthanized.

After sacrifice each condyle was separated from the tibial plateau and all soft tissue was removed with a scalpel. Joints were dehydrated using increasing concentrations of ethanol and embedded in polymethylmethacrylate (PMMA) using a published procedure51,52.

μCT imaging

Prior to histology and histomorphometry, a μCT machine with 18μm resolution was used to scan 17mm diameter bone cores, which contained either the biomimetic or the simple porous scaffolds from the lateral condyles. Scaffolds implanted in the medial condyles contained strain gauges that caused artifacts and prevented μCT measurement collection. A region of interest corresponding to the scaffold region analyzed using histomorphometry (Fig regions) was analyzed using a threshold value of 200. The Scanco software was used to measure bone volume (BV/TV, %), trabecular number (mm-1), trabecular spacing (mm), and trabecular thickness (mm).

Histology and Histomorphometry

After the scaffolds were embedded in PMMA and scanned with the μCT, four to six sagittal sections were taken through every scaffold using a Bronwell diamond waffering saw (Rochester, NY). The slides were ground and polished to 100μm using a Leco grinder and polisher (St. Joseph, MI). Each slide was stained with Mineralized Bone Stain (Harrington Arthritis Research Center, Phoenix, AZ). Three sections through the center of each scaffold were analyzed using histology and histomorphometry.

An Olympus MagnaFire SP digital camera (Olympus, Tokyo, Japan) coupled to a Nikon Optiphot microscope (Nikon, Tokyo, Japan) was used to collect images of all the slides. Each slide was broken up into 4 different regions according to a previously published procedure28. The four regions were the deep periscaffold, periscaffold, scaffold, and the superficial scaffold region (Figure 6). The scaffold region was composed of all implant materials including PBT and any attached epoxy or strain gauges in the wired scaffolds. The periscaffold region was 1mm to either side of the scaffold. The deep periscaffold region was 1mm deep to the scaffold. The superficial scaffold region was the area superficial to the scaffold and was the area that articulated with the adjacent tibial plateau in vivo (Figure 6). These regions were defined to help normalize the changes in bone volume that occur in these locations as seen in previous studies28.

Figure 6.

Figure 6

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Schematic diagram depicting the regions of the scaffold (deep periscaffold, periscaffold, scaffold, and superficial) which were analyzed using histomorphometry.

In each region bone perimeter and area (mm2), osteoid area (mm2), marrow area (mm2), and synthetic material area (mm2) were measured using light microscopy. Fluorescent light was used to measure single and double label length (mm) as well as interlabel distance (mm). All measurements were performed using Image J (NIH, Bethesda, MD) at 62.5x. Histological and histomorphometric calculations were performed according to the guidelines set forth by the American Society for Bone and Mineral Research53. Bone, Osteoid, and Marrow volume were calculated as a percent of total pore volume to normalize for any differences in porosity between scaffolds. The percent labeled surface was calculated as half of the single label + the average of the inner and outer double labels.

All parameters were compared between simple porous and biomimetic scaffold with ANOVA using SPSS. The threshold for statistical significance was set at p≤0.05. All data is reported as an average ± standard deviation.

Results

All eight test animals were weight bearing within a day of surgery and returned to a full running regimen two weeks post-op. One simple porous scaffold was excluded from the study because medial and lateral scaffold placements were performed during separate surgeries, and the lateral scaffold was only in place for 1.5 months. For this reason eight biomimetic scaffolds and seven simple porous scaffolds were compared. μCT analysis of scaffold porosity prior to implantation showed that the simple porous scaffolds had an average porosity of 48.7±2.33% while the biomimetic scaffolds had a porosity of 48.6±.84%. Average stiffness was 1.53±.17GPa for the simple porous structure and 1.46±.34GPa for the biomimetic trabecular implant. The trabecular bone stiffness in the area where the scaffolds were placed (distal femoral condyles) has been reported in the range of 0.25 to 0.36GPa54.

Histology showed a significant (p=.009) 506% increase in bone ingrowth into the scaffold region of the biomimetic scaffold design compared to the simple porous scaffold (Figure 7). Additionally, in the periscaffold and deep periscaffold regions there was a 40% and 95% (p=.005 and .028) increase in bone volume respectively (Figure 7). There was a 66% and 70% (p=.002 and .034) decrease in soft tissue found within the periscaffold and deep periscaffold region (Table 1), and a 22% (p=.031) decrease in tetracycline labeled surface within the periscaffold region (Table 1).

Figure 7.

Figure 7

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(A) and (B) are images representing overall scaffold location and bone growth for the simple porous (A) and biomimetic (B) scaffolds. Magnified images showing increased bone growth into and along the lateral edges of the biomimetic scaffold (D) and (F) compared to the simple porous scaffold (C) and (E). Within each image (B) indicates bone and (P) indicates PBT scaffold material. All scale bars are 1mm.

Table 1.

Histomorphometry results

Simple Biomimetic p Values
Scaffold
BV/TV (%) 2.00±2.44 12.15±8.34 0.009
MaV/TV (%) 44.78±20.94 40.80±16.17 0.685
OsV/TV (%) 53.21±21.77 47.05±18.23 0.561
Lbl sur (%) 1.98±2.15 3.41±1.78 0.182
Periscaffold
BV/TV (%) 41.39±11.28 58.13±7.59 0.005
MaV/TV (%) 30.29±14.14 32.19±7.87 0.75
OsV/TV (%) 28.32±13.68 9.68±3.07 0.002
Lbl sur (%) 6.28±1.09 4.87±1.15 0.031
Deep
BV/TV (%) 15.74±12.12 30.67±11.19 0.028
MaV/TV (%) 39.58±27.05 56.00±12.63 0.147
OsV/TV (%) 44.68±35.80 13.32±10.94 0.034
Lbl sur (%) 4.37±2.24 5.55±2.72 0.382
Superficial
BV/TV (%) 17.52±9.18 13.07±4.86 0.253
MaV/TV (%) 6.83±2.15 6.76±5.09 0.971
OsV/TV (%) 75.64±11.21 80.17±6.07 0.339
Lbl sur (%) 4.08±2.23 3.05±1.34 0.288
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BV = Bone volume

MaV = Marrow Volume

OsV = Osteoid Volume

TV = Total porous volume (BV + MaV + OsV)

Lbl sur = Labeled surface

μCT analysis showed a significant (p=.029) 629% increase in bone volume within the scaffold region of trabeculated implants (Figure 8). Additionally, there was a 454% (p=.018) increase in trabecular number and an 82% (p=.045) decrease in trabecular spacing within the scaffold region. Trabecular thickness was similar in the two scaffold types (p=.089) (Table 2).

Figure 8.

Figure 8

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CT images showing the markedly increased bone growth within the biomimetic scaffold (B) and (D) compared to the simple porous scaffold (A) and (C). Bone growth throughout the entire scaffold is depicted in (A) and (B), while (C) and (D) are cross sections taken through the center of the scaffolds

Table 2.

MicroCT results

Simple Biomimetic p Values
BV/TV (%) 1.44±.47 10.5±4.56 0.029
Trab Number (1/mm) .11±.03 .61±.21 0.018
Trab Thickness (mm) .12±.01 .17±.04 0.089
Trab Spacing (mm) 9.31±2.88 1.63±.62 0.045
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BV/TV is bone volume as a percent of total volume (Bone, Marrow, and Osteoid)

Discussion

PBT scaffolds offer a number of advantages for their use as tissue engineering scaffolds: they are able to withstand physiological loads when placed in weight bearing locations, they have a slow resorption rate which allows more time for bone ingrowth prior to polymer degradation and they can be easily manufactured with highly controlled structures. The high degree of control over polymer scaffolds created using solid free form fabrication allows the design of biomimetic internal architectures unique to each patient and defect site. It is important to determine whether a biomimetic scaffold accelerates bone ingrowth compared to simple porous designs because use of complex custom implants increases the cost of scaffold production and necessitates patient exposure to additional radiation during μCT imaging. This is the first study we know of that has compared bone ingrowth into macroscopic porous scaffolds, which had either a simple pore structure or a complex pore structure mimicking the native tissue architecture. Previous studies have looked at the effect of microscopic biomimetic surface modifications on bone growth55,56 but not macroscopic biomimetic structures. Studies on modification of macroscopic structure have analyzed the effects of pore size, pore interconnectivity and pore shape on bone growth, but these studies use synthetic shapes instead of replicating the native architecture32-37,57.

Both μCT analysis and histology showed more than a 500% increase in bone volume within biomimetic porous scaffolds compared to the simple porous scaffolds, indicating that scaffold architectures mimicking native bone tissue facilitate faster bone growth. While the biomimetic scaffold had a higher bone volume within the scaffold pores, both scaffold types had a similar amount of tissue volume (bone plus osteoid volumes). Similar amounts of tissue within scaffold pores indicates that cells and tissue had similar access to the internal porous spaces within the scaffolds; however, the increased bone volume in the biomimetic implants indicates that the architecture facilitated accelerated bone formation. The μCT analysis showed no significant difference between the trabecular bone thickness within the biomimetic and simple porous scaffolds. This suggests that the structure of individual trabeculae within both types of scaffolds is similar and that altering scaffold architecture changes the amount of bone as seen by the increase in trabecular number, but not its microscopic structure.

Although the largest increase in bone volume was seen within the scaffold pores, histology also showed an increase in bone volume deep and adjacent to the biomimetic scaffolds. This indicates that the effects of the biomimetic scaffolds extended beyond the porous region to areas adjacent to the scaffold. The increased bone volume and decreased osteoid volume in the areas adjacent to the biomimetic scaffolds suggests that the tissue surrounding these implants is more completely mineralized at this time point. This is further supported by the decreased percentage of labeled surface in the periscaffold region of the trabeculated implants.

Creation of a PBT implant with a simple interconnected porous pattern yields a scaffold that does not have sufficient mechanical integrity to be implanted into a load-bearing defect until a solid outer cylinder is built around the porous pattern to improve strength. Bone ingrowth windows were incorporated into this simple porous scaffold design to allow tissue growth into the scaffolds from the lateral surfaces of the implant. Use of large square windows was not possible with the simple porous design as mechanical loading damaged parts of the scaffold in unsupported areas. The internal architecture of biomimetic trabeculated scaffolds allowed use of larger square windows instead of the smaller round windows used in the simple porous scaffolds. This was because there were fewer points of contact between the containment ring but greater internal connectivity in the trabeculated design allowing further erosion of the outer cylinder without compromising the ability of the scaffold to withstand mechanical loading. It is possible that the changes in window size and shape contributed to the accelerated bone growth in the biomimetic scaffolds.

One possible explanation for how window size alters bone formation is that decreased window size may reduce cell and tissue access to the interior porous regions of the scaffolds, potentially contributing to decreased bone formation in the simple porous scaffold. However, both scaffolds had the same “tissue” volume (bone plus osteoid volumes) suggesting that the window shape and size did not influence cell and tissue access to the scaffold pores in vivo. Another possible explanation for the effect that window size and shape has on bone growth is that a smaller opening may decrease vessel growth into the scaffold pores, which could result in decreased de novo bone formation because of decreased osteogenic precursor cells and other osteogenic factors. This warrants further investigation in future studies. Additionally, future studies using different materials may be able to look at pore designs without having to include an outer skin for mechanical stability and this would eliminate any effect that windows play on bone formation.

Additionally, there was increased bone growth evident deep and lateral to the trabeculated scaffolds. Differences in the size and shape of bone ingrowth windows would not have directly affected the areas adjacent to the scaffold because they do not limit tissue access to these areas. This suggests that the biomimetic scaffold has effects extending beyond the confines of the implant. This may be due to differences in the loading environment resulting from the different implant designs. While the overall stiffness was similar for the two scaffold designs, the differences in load transmission to the surrounding implant sites were not studied and might account for the variances in bone formation within and adjacent to the scaffolds.

Previous studies using hydroxyapatite, which has some biomimetic characteristics, have shown good osteoconductivity at the implant site demonstrating that replicating the native environment encourages faster bone growth58. An increased osteoblastic phenotype and more rapid mineralization have also been noted when testing biomimetic surface morphology55,59,60 in cell culture and in vivo. These studies have not focused on the effects of biomimetic macroporous architectures, but instead focus on altering the surface morphology of implants using simulated body fluid treatments55 and thermal phase separation59,60. The results of these previous studies suggest that mimicking native bone tissue increases bone growth. This is supported by our results, which add to previous studies by showing that biomimetic macroscopic porous architecture also accelerates bone formation.

Accelerating bone ingrowth into scaffolds is advantageous because sufficient bone growth is needed prior to polymer degradation to provide tissue engineered constructs and regenerating tissues with continuous support. Ideally, stabilizing bone growth needs to occur before polymer degradation, as degradation prior to bone formation can result in fibrous tissue formation that will not support loads in weight bearing locations. Various techniques have been utilized to maximize bone growth into polymer scaffolds including the use of osteoconductive calcium phosphate ceramic particles and osteoinductive proteins27-29. The results of this study suggest that altering scaffold architecture, when utilized in conjunction with these other treatments, promotes even faster bone growth and may provide a polymer construct suitable to treating large bone defects as an alternative to using allografts and autografts. Additionally, increased bone growth which improves scaffold stability may enhance integration of tissue engineered cartilage with native cartilage tissue28. The accelerated stabilization would also be advantageous in other tissue engineering applications at musculoskeletal interfaces such as ligamentous or tendon insertions.

In addition to the clinical implications of faster bone to implant bonding, the accelerated bonding allows more rapid collection of accurate load measurements from the strain-gauged scaffolds27. These sensate scaffolds are currently the only available system to directly measure native joint loads. Previous in vivo studies utilizing sensate scaffolds27 have shown that implant loads increase as a function of time post-op while bone ingrowth and healing occur. When bone ingrowth is complete, the sensate implants provide accurate measurement of load changes during routine weight bearing activities27,28, joint loading following surgical resection of the ACL to induce osteoarthritis61, and changes in joint loading after administration of various pharmacological treatments. Additionally, previous studies have shown that replicating in vivo conditions, i.e. physiological joint loading environments, can be utilized in tissue engineering applications to improve engineered tissue quality and progress toward a functional construct62-67.

This study used histology and μCT to demonstrate that a biomimetic trabeculated porous architecture accelerates bone growth into and around PBT polymer implants. This results in quicker stabilization of the scaffolds which will likely lead to increased success rates for tissue engineered constructs and better joint load monitoring. The findings of this study suggest that the additional cost and exposure to radiation may be worthwhile to create biomimetic implants that accelerate bone formation. Studies to determine whether exact patient specific trabeculated designs offer an advantage over a generic trabeculated design will be needed to further address the question of whether biomimetic designs unique to each patient and defect site further accelerate healing.

Acknowledgements

The authors thank the NIH-NIBIB for support through grant RO1 - EB000660.

References

  • 1.Muscolo DL, Ayerza MA, Aponte-Tinao LA, Ranalletta M. Use of distal femoral osteoarticular allografts in limb salvage surgery. Surgical technique. J Bone Joint Surg Am. 2006;88(Suppl 1 Pt 2):305–21. doi: 10.2106/jbjs.f.00324. [DOI] [PubMed] [Google Scholar]
  • 2.Sen MK, Miclau T. Autologous iliac crest bone graft: should it still be the gold standard for treating nonunions? Injury. 2007;38(Suppl 1):S75–80. doi: 10.1016/j.injury.2007.02.012. [DOI] [PubMed] [Google Scholar]
  • 3.Ayerza MA, Aponte-Tinao LA, Abalo E, Muscolo DL. Continuity and function of patellar tendon host-donor suture in tibial allograft. Clin Orthop Relat Res. 2006;450:33–8. doi: 10.1097/01.blo.0000229291.21722.b5. [DOI] [PubMed] [Google Scholar]
  • 4.Johari A, Shingade V, Lobo Gajiwala A, Shah V, C DL. The use of irradiated allograft in a paediatric population: an Indian experience. Cell Tissue Bank. 2007;8(1):13–22. doi: 10.1007/s10561-006-9001-4. [DOI] [PubMed] [Google Scholar]
  • 5.Rajan GP, Fornaro J, Trentz O, Zellweger R. Cancellous allograft versus autologous bone grafting for repair of comminuted distal radius fractures: a prospective, randomized trial. J Trauma. 2006;60(6):1322–9. doi: 10.1097/01.ta.0000195977.18035.40. [DOI] [PubMed] [Google Scholar]
  • 6.Kessler P, Thorwarth M, Bloch-Birkholz A, Nkenke E, Neukam FW. Harvesting of bone from the iliac crest--comparison of the anterior and posterior sites. Br J Oral Maxillofac Surg. 2005;43(1):51–6. doi: 10.1016/j.bjoms.2004.08.026. [DOI] [PubMed] [Google Scholar]
  • 7.Pacifico MD, Floyd D, Wood SH. Tibial stress fracture as a complication of free-fibula vascularised graft for mandibular reconstruction. Br J Plast Surg. 2003;56(8):832–4. doi: 10.1016/j.bjps.2003.08.004. [DOI] [PubMed] [Google Scholar]
  • 8.Heary RF, Schlenk RP, Sacchieri TA, Barone D, Brotea C. Persistent iliac crest donor site pain: independent outcome assessment. Neurosurgery. 2002;50(3):510–6. doi: 10.1097/00006123-200203000-00015. discussion 516-7. [DOI] [PubMed] [Google Scholar]
  • 9.Silber JS, Anderson DG, Daffner SD, Brislin BT, Leland JM, Hilibrand AS, Vaccaro AR, Albert TJ. Donor site morbidity after anterior iliac crest bone harvest for single-level anterior cervical discectomy and fusion. Spine. 2003;28(2):134–9. doi: 10.1097/00007632-200301150-00008. [DOI] [PubMed] [Google Scholar]
  • 10.Zimmermann CE, Borner BI, Hasse A, Sieg P. Donor site morbidity after microvascular fibula transfer. Clin Oral Investig. 2001;5(4):214–9. doi: 10.1007/s00784-001-0140-5. [DOI] [PubMed] [Google Scholar]
  • 11.Buck BE, Malinin TI, Brown MD. Bone transplantation and human immunodeficiency virus. An estimate of risk of acquired immunodeficiency syndrome (AIDS) Clin Orthop Relat Res. 1989;240:129–36. [PubMed] [Google Scholar]
  • 12.Lewandrowski KU, Rebmann V, Passler M, Schollmeier G, Ekkernkamp A, Grosse-Wilde H, Tomford WW. Immune response to perforated and partially demineralized bone allografts. J Orthop Sci. 2001;6(6):545–55. doi: 10.1007/s007760100011. [DOI] [PubMed] [Google Scholar]
  • 13.Pilliar RM. Porous-surfaced metallic implants for orthopedic applications. J Biomed Mater Res. 1987;21(A1 Suppl):1–33. [PubMed] [Google Scholar]
  • 14.Bush JL, Wilson JB, Vail TP. Management of bone loss in revision total knee arthroplasty. Clin Orthop Relat Res. 2006;452:186–92. doi: 10.1097/01.blo.0000229360.04620.93. [DOI] [PubMed] [Google Scholar]
  • 15.Kirkpatrick JS, Venugopalan R, Beck P, Lemons J. Corrosion on spinal implants. J Spinal Disord Tech. 2005;18(3):247–51. [PubMed] [Google Scholar]
  • 16.Goodman SB. Wear particles, periprosthetic osteolysis and the immune system. Biomaterials. 2007;28(34):5044–8. doi: 10.1016/j.biomaterials.2007.06.035. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 17.Parsons JR. Resorbable materials and composites. New concepts in orthopedic biomaterials. Orthopedics. 1985;8(7):907–15. doi: 10.3928/0147-7447-19850701-17. [DOI] [PubMed] [Google Scholar]
  • 18.Claase MB, Grijpma DW, Mendes SC, De Bruijn JD, Feijen J. Porous PEOT/PBT scaffolds for bone tissue engineering: preparation, characterization, and in vitro bone marrow cell culturing. J Biomed Mater Res A. 2003;64(2):291–300. doi: 10.1002/jbm.a.10418. [DOI] [PubMed] [Google Scholar]
  • 19.Liu Q, de Wijn JR, van Blitterswijk CA. Composite biomaterials with chemical bonding between hydroxyapatite filler particles and PEG/PBT copolymer matrix. J Biomed Mater Res. 1998;40(3):490–7. doi: 10.1002/(sici)1097-4636(19980605)40:3<490::aid-jbm20>3.0.co;2-m. [DOI] [PubMed] [Google Scholar]
  • 20.Radder AM, Leenders H, van Blitterswijk CA. Interface reactions to PEO/PBT copolymers (Polyactive) after implantation in cortical bone. J Biomed Mater Res. 1994;28(2):141–51. doi: 10.1002/jbm.820280202. [DOI] [PubMed] [Google Scholar]
  • 21.Radder AM, Leenders H, van Blitterswijk CA. Application of porous PEO/PBT copolymers for bone replacement. J Biomed Mater Res. 1996;30(3):341–51. doi: 10.1002/(SICI)1097-4636(199603)30:3<341::AID-JBM8>3.0.CO;2-Q. [DOI] [PubMed] [Google Scholar]
  • 22.Risbud M, Saheb DN, Jog J, Bhonde R. Preparation, characterization and in vitro biocompatibility evaluation of poly(butylene terephthalate)/wollastonite composites. Biomaterials. 2001;22(12):1591–7. doi: 10.1016/s0142-9612(00)00325-2. [DOI] [PubMed] [Google Scholar]
  • 23.Sakkers RJ, Dalmeyer RA, de Wijn JR, van Blitterswijk CA. Use of bone-bonding hydrogel copolymers in bone: an in vitro and in vivo study of expanding PEO-PBT copolymers in goat femora. J Biomed Mater Res. 2000;49(3):312–8. doi: 10.1002/(sici)1097-4636(20000305)49:3<312::aid-jbm3>3.0.co;2-e. [DOI] [PubMed] [Google Scholar]
  • 24.Sakkers RJ, de Wijn JR, Dalmeyer RA, van Blitterswijk CA, Brand R. Evaluation of copolymers of polyethylene oxide and polybutylene terephthalate (polyactive): mechanical behaviour. J Mater Sci Mater Med. 1998;9(7):375–9. doi: 10.1023/a:1013227428530. [DOI] [PubMed] [Google Scholar]
  • 25.Woodfield TB, Bezemer JM, Pieper JS, van Blitterswijk CA, Riesle J. Scaffolds for tissue engineering of cartilage. Crit Rev Eukaryot Gene Expr. 2002;12(3):209–36. doi: 10.1615/critreveukaryotgeneexpr.v12.i3.40. [DOI] [PubMed] [Google Scholar]
  • 26.Roessler M, Wilke A, Griss P, Kienapfel H. Missing osteoconductive effect of a resorbable PEO/PBT copolymer in human bone defects: a clinically relevant pilot study with contrary results to previous animal studies. J Biomed Mater Res. 2000;53(2):167–73. doi: 10.1002/(sici)1097-4636(2000)53:2<167::aid-jbm6>3.0.co;2-j. [DOI] [PubMed] [Google Scholar]
  • 27.Geffre CP, Bliss CL, Szivek JA, Deyoung DW, Ruth JT, Margolis DS. Sensate scaffolds coupled to telemetry can monitor in vivo loading from within a joint over extended periods of time. J Biomed Mater Res B Appl Biomater. 2008;84(1):263–70. doi: 10.1002/jbm.b.30869. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 28.Szivek JA, Bliss CL, Geffre CP, Margolis DS, DeYoung DW, Ruth JT, Schnepp AB, Tellis BC, Vaidyanathan RK. An instrumented scaffold can monitor loading in the knee joint. J Biomed Mater Res B Appl Biomater. 2006;79(2):218–28. doi: 10.1002/jbm.b.30532. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 29.Szivek JA, Margolis DS, Garrison BK, Nelson E, Vaidyanathan RK, DeYoung DW. TGF-beta1-enhanced TCP-coated sensate scaffolds can detect bone bonding. J Biomed Mater Res B Appl Biomater. 2005;73(1):43–53. doi: 10.1002/jbm.b.30177. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 30.Szivek JA, Roberto RF, Slack JM, Majeed BS. An implantable strain measurement system designed to detect spine fusion: preliminary results from a biomechanical in vivo study. Spine. 2002;27(5):487–97. doi: 10.1097/00007632-200203010-00009. [DOI] [PubMed] [Google Scholar]
  • 31.Bliss CL, Szivek JA, Tellis BC, Margolis DS, Schnepp AB, Ruth JT. Sensate scaffolds can reliably detect joint loading. J Biomed Mater Res B Appl Biomater. 2007;81(1):30–9. doi: 10.1002/jbm.b.30632. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 32.Claase MB, de Bruijn JD, Grijpma DW, Feijen J. Ectopic bone formation in cell-seeded poly(ethylene oxide)/poly(butylene terephthalate) copolymer scaffolds of varying porosity. J Mater Sci Mater Med. 2007;18(7):1299–307. doi: 10.1007/s10856-006-0077-y. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 33.Goodman S, Toksvig-Larsen S, Aspenberg P. Ingrowth of bone into pores in titanium chambers implanted in rabbits: effect of pore cross-sectional shape in the presence of dynamic shear. J Biomed Mater Res. 1993;27(2):247–53. doi: 10.1002/jbm.820270215. [DOI] [PubMed] [Google Scholar]
  • 34.Hing KA, Best SM, Tanner KE, Bonfield W, Revell PA. Mediation of bone ingrowth in porous hydroxyapatite bone graft substitutes. J Biomed Mater Res A. 2004;68(1):187–200. doi: 10.1002/jbm.a.10050. [DOI] [PubMed] [Google Scholar]
  • 35.Jones AC, Arns CH, Sheppard AP, Hutmacher DW, Milthorpe BK, Knackstedt MA. Assessment of bone ingrowth into porous biomaterials using MICRO-CT. Biomaterials. 2007;28(15):2491–504. doi: 10.1016/j.biomaterials.2007.01.046. [DOI] [PubMed] [Google Scholar]
  • 36.Klawitter JJ, Bagwell JG, Weinstein AM, Sauer BW. An evaluation of bone growth into porous high density polyethylene. J Biomed Mater Res. 1976;10(2):311–23. doi: 10.1002/jbm.820100212. [DOI] [PubMed] [Google Scholar]
  • 37.Mastrogiacomo M, Scaglione S, Martinetti R, Dolcini L, Beltrame F, Cancedda R, Quarto R. Role of scaffold internal structure on in vivo bone formation in macroporous calcium phosphate bioceramics. Biomaterials. 2006;27(17):3230–7. doi: 10.1016/j.biomaterials.2006.01.031. [DOI] [PubMed] [Google Scholar]
  • 38.Bobyn JD, Pilliar RM, Cameron HU, Weatherly GC. The optimum pore size for the fixation of porous-surfaced metal implants by the ingrowth of bone. Clin Orthop Relat Res. 1980;150:263–70. [PubMed] [Google Scholar]
  • 39.Flautre B, Descamps M, Delecourt C, Blary MC, Hardouin P. Porous HA ceramic for bone replacement: role of the pores and interconnections - experimental study in the rabbit. J Mater Sci Mater Med. 2001;12(8):679–82. doi: 10.1023/a:1011256107282. [DOI] [PubMed] [Google Scholar]
  • 40.Oh SH, Park IK, Kim JM, Lee JH. In vitro and in vivo characteristics of PCL scaffolds with pore size gradient fabricated by a centrifugation method. Biomaterials. 2007;28(9):1664–71. doi: 10.1016/j.biomaterials.2006.11.024. [DOI] [PubMed] [Google Scholar]
  • 41.Robinson BP, Hollinger JO, Szachowicz EH, Brekke J. Calvarial bone repair with porous D,L-polylactide. Otolaryngol Head Neck Surg. 1995;112(6):707–13. doi: 10.1016/S0194-59989570180-X. [DOI] [PubMed] [Google Scholar]
  • 42.Hollister SJ, Levy RA, Chu TM, Halloran JW, Feinberg SE. An image-based approach for designing and manufacturing craniofacial scaffolds. Int J Oral Maxillofac Surg. 2000;29(1):67–71. doi: 10.1034/j.1399-0020.2000.290115.x. [DOI] [PubMed] [Google Scholar]
  • 43.Hollister SJ, Maddox RD, Taboas JM. Optimal design and fabrication of scaffolds to mimic tissue properties and satisfy biological constraints. Biomaterials. 2002;23(20):4095–103. doi: 10.1016/s0142-9612(02)00148-5. [DOI] [PubMed] [Google Scholar]
  • 44.Shi T, Yue X, Xiong Z, Zhang R, Yan Y. Rapid manufacturing of degradable porous polymer bone scaffold. Sheng Wu Yi Xue Gong Cheng Xue Za Zhi. 2002;19(2):348–9. [PubMed] [Google Scholar]
  • 45.Tellis BC, Szivek JA, Bliss CL, Margolis DS, Vaidyanathan RK, Calvert P. Trabecular scaffolds created using micro CT guided fused deposition modeling. J of Materials Science & Engineering-C. 2008;28(1):171–8. doi: 10.1016/j.msec.2006.11.010. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 46.Cordaro NM, Szivek JA, DeYoung DW. Surface enhancements accelerate bone bonding to CPC-coated strain gauges. J Biomed Mater Res. 2001;56(1):109–19. doi: 10.1002/1097-4636(200107)56:1<109::aid-jbm1075>3.0.co;2-w. [DOI] [PubMed] [Google Scholar]
  • 47.Rabkin BA, Szivek JA, Schonfeld JE, Halloran BP. Long-term measurement of bone strain in vivo: the rat tibia. J Biomed Mater Res. 2001;58(3):277–81. doi: 10.1002/1097-4636(2001)58:3<277::aid-jbm1017>3.0.co;2-t. [DOI] [PubMed] [Google Scholar]
  • 48.Szivek JA, Anderson PL, DeYoung DW. In vivo strain measurements collected using calcium phosphate ceramic-bonded strain gauges. J Invest Surg. 1997;10(5):263–73. doi: 10.3109/08941939709032165. [DOI] [PubMed] [Google Scholar]
  • 49.Battraw GA, Szivek JA, Anderson PL. Interface strength studies of calcium phosphate ceramic coated strain gauges. J Biomed Mater Res. 1998;43(4):462–8. doi: 10.1002/(sici)1097-4636(199824)43:4<462::aid-jbm14>3.0.co;2-j. [DOI] [PubMed] [Google Scholar]
  • 50.Szivek JA, Nelson ER, Hajdu SD, Yablonski K, DeYoung DW. Transforming growth factor-beta1 accelerates bone bonding to a blended calcium phosphate ceramic coating: a dose-response study. J Biomed Mater Res A. 2004;68(3):537–43. doi: 10.1002/jbm.a.20099. [DOI] [PubMed] [Google Scholar]
  • 51.Emmanual J, Hornbeck C, Bloebaum RD. A polymethyl methacrylate method for large specimens of mineralized bone with implants. Stain Technol. 1987;62(6):401–10. doi: 10.3109/10520298709108030. [DOI] [PubMed] [Google Scholar]
  • 52.Martini F, Leichtle U, Lebherz C. Optimized polymethyl methacrylate-embedding enables exact section examination of the femur with an uncemented femoral stem. Z Orthop Ihre Grenzgeb. 2001;139(6):531–5. doi: 10.1055/s-2001-19236. [DOI] [PubMed] [Google Scholar]
  • 53.Parfitt AM. Bone histomorphometry: proposed system for standardization of nomenclature, symbols, and units. Calcif Tissue Int. 1988;42(5):284–6. doi: 10.1007/BF02556360. [DOI] [PubMed] [Google Scholar]
  • 54.Kuhn JL, Goldstein SA, Ciarelli MJ, Matthews LS. The limitations of canine trabecular bone as a model for human: a biomechanical study. J Biomech. 1989;22(2):95–107. doi: 10.1016/0021-9290(89)90032-8. [DOI] [PubMed] [Google Scholar]
  • 55.Fan H, Ikoma T, Tanaka J, Zhang X. Surface structural biomimetics and the osteoinduction of calcium phosphate biomaterials. J Nanosci Nanotechnol. 2007;7(3):808–13. doi: 10.1166/jnn.2007.501. [DOI] [PubMed] [Google Scholar]
  • 56.Tuzlakoglu K, Bolgen N, Salgado AJ, Gomes ME, Piskin E, Reis RL. Nano- and microfiber combined scaffolds: a new architecture for bone tissue engineering. J Mater Sci Mater Med. 2005;16(12):1099–104. doi: 10.1007/s10856-005-4713-8. [DOI] [PubMed] [Google Scholar]
  • 57.Simon JL, Michna S, Lewis JA, Rekow ED, Thompson VP, Smay JE, Yampolsky A, Parsons JR, Ricci JL. In vivo bone response to 3D periodic hydroxyapatite scaffolds assembled by direct ink writing. J Biomed Mater Res A. 2007;83(3):747–58. doi: 10.1002/jbm.a.31329. [DOI] [PubMed] [Google Scholar]
  • 58.Wang H, Li Y, Zuo Y, Li J, Ma S, Cheng L. Biocompatibility and osteogenesis of biomimetic nano-hydroxyapatite/polyamide composite scaffolds for bone tissue engineering. Biomaterials. 2007;28(22):3338–48. doi: 10.1016/j.biomaterials.2007.04.014. [DOI] [PubMed] [Google Scholar]
  • 59.Woo KM, Jun JH, Chen VJ, Seo J, Baek JH, Ryoo HM, Kim GS, Somerman MJ, Ma PX. Nano-fibrous scaffolding promotes osteoblast differentiation and biomineralization. Biomaterials. 2007;28(2):335–43. doi: 10.1016/j.biomaterials.2006.06.013. [DOI] [PubMed] [Google Scholar]
  • 60.Chen VJ, Smith LA, Ma PX. Bone regeneration on computer-designed nano-fibrous scaffolds. Biomaterials. 2006;27(21):3973–9. doi: 10.1016/j.biomaterials.2006.02.043. [DOI] [PubMed] [Google Scholar]
  • 61.Szivek JA, Geffre CP, Margolis DS, Deyoung DW, Ruth JT, Harris DT. A Cartilage Covered “Sensate” Scaffold Can be Used to Measure Knee Load Changes In Vivo Following ACL Disruption. 6th Combined Meeting of the Orthopaedic Research Societies; Honolulu, Hawaii. 2007 October. [Google Scholar]
  • 62.Lee DA, Bader DL. Compressive strains at physiological frequencies influence the metabolism of chondrocytes seeded in agarose. J Orthop Res. 1997;15(2):181–8. doi: 10.1002/jor.1100150205. [DOI] [PubMed] [Google Scholar]
  • 63.Neidlinger-Wilke C, Grood ES, Wang J-C, Brand RA, Claes L. Cell alignment is induced by cyclic changes in cell length: studies of cells grown in cyclically stretched substrates. J Orthop Res. 2001;19(2):286–93. doi: 10.1016/S0736-0266(00)00029-2. [DOI] [PubMed] [Google Scholar]
  • 64.Smith RL, Lin J, Trindade MC, Shida J, Kajiyama G, Vu T, Hoffman AR, van der Meulen MC, Goodman SB, Schurman DJ, et al. Time-dependent effects of intermittent hydrostatic pressure on articular chondrocyte type II collagen and aggrecan mRNA expression. J Rehabil Res Dev. 2000;37(2):153–61. [PubMed] [Google Scholar]
  • 65.Waldman SD, Couto DC, Grynpas MD, Pilliar RM, Kandel RA. A single application of cyclic loading can accelerate matrix deposition and enhance the properties of tissue-engineered cartilage. Osteoarthritis Cartilage. 2006;14(4):323–30. doi: 10.1016/j.joca.2005.10.007. [DOI] [PubMed] [Google Scholar]
  • 66.Waldman SD, Spiteri CG, Grynpas MD, Pilliar RM, Hong J, Kandel RA. Effect of biomechanical conditioning on cartilaginous tissue formation in vitro. J Bone Joint Surg Am. 2003;85-A(Suppl 2):101–5. doi: 10.2106/00004623-200300002-00013. [DOI] [PubMed] [Google Scholar]
  • 67.Waldman SD, Spiteri CG, Grynpas MD, Pilliar RM, Kandel RA. Long-term intermittent compressive stimulation improves the composition and mechanical properties of tissue-engineered cartilage. Tissue Eng. 2004;10(910):1323–31. doi: 10.1089/ten.2004.10.1633. [DOI] [PubMed] [Google Scholar]

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