In vivo evaluation of the regenerative capability of glycylglycine ethyl ester-substituted polyphosphazene and poly (lactic-co-glycolic acid) blends: A rabbit critical-sized bone defect model

Abstract

In an effort to understand the biological capability of polyphosphazene-based polymers, three-dimensional (3D) biomimetic bone scaffolds were fabricated using blends of poly[(glycine ethylglycinato)₇₅(phenylphenoxy)₂₅]phosphazene (PNGEGPhPh) and poly(lactic-co-glycolic acid) (PLGA), and an in vivo evaluation was performed in a rabbit critical-sized bone defect model. The matrices constructed from PNGEGPhPh-PLGA blends were surgically implanted into 15mm critical-sized radial defects of the rabbits as structural templates for bone tissue regeneration. PLGA, which is the most commonly used synthetic bone graft substitute, was used as a control in this study. Radiological and histological analyses demonstrated that PNGEGPhPh-PLGA blends exhibited favorable in vivo biocompatibility and osteoconductivity as the newly designed matrices allowed new bone formation to occur without adverse immunoreactions. X-ray images of the blends showed higher levels of radiodensity than that of the pristine PLGA, indicating higher rates of new bone formation and regeneration. Micro-computed tomography instrument (micro-CT) quantification revealed that new bone volume fractions were significantly higher for the PNGEGPhPh-PLGA blends than for the PLGA controls after 4 weeks. The new bone volume increased linearly with increasing time points, with the new tissues observed throughout the defect area for the blend and only at the implant site’s extremes for the PLGA control. Histologically, the polyphosphazene system appeared to show tissue responses and bone ingrowths superior to that of PLGA. By the end of the study, the defects with PNGEGPhPh-PLGA scaffolds exhibited evidence of effective bone tissue-ingrowth and minimal inflammatory responses. Thus, polyphosphazene-containing biomaterials have excellent translational potential for use in bone regenerative engineering applications.

Graphical Abstarct

INTRODUCTION

Polymers have played essential and indispensable roles in wide-ranging clinical applications, including controlled drug delivery, sutures, and vascular grafts, among others.^1-3 In orthopaedic surgery, biodegradable polymers have become increasingly important in their utilization as scaffolding materials for bone tissue repair and regeneration.^4-6 The repair of bone defects has been based on two grafting options, namely, autografts and allografts⁷. Autogenous bone grafts are bone tissues transplanted from one site to another of the same individual’s body, and they are currently considered as the gold standard for most surgical reconstruction procedures.⁷ Autografts possess benefits, such as optimal biological behavior, histocompatibility, and no risk of disease transmission. However, they suffer from drawbacks such as donor-site morbidity, pain, increased operative blood loss, and unavailability of large tissue volumes.¹ Allografts, on the other hand, are tissues taken from one person (donor) to another person (recipient), and these bone grafts are often obtained from cadavers.^{1, 8} Donor-site morbidity is not an issue for allogenic bone grafts; however, some of the associated limitations include immune rejection, risk of communicable diseases, and donor availability.¹ These drawbacks currently being faced by autografts and allografts have stimulated the desire for the design of synthetic alternatives that maximize the benefits and mitigate the disadvantages of the biological bone grafts.^{1, 7}

Recent advances in materials science and engineering have laid the groundwork for the design of optimal polymeric materials with unique biological recognition properties.^9-11 These materials have the ability to regulate and coordinate cellular activities toward functional tissue development.^12-15 Consequently, the syntheses of a variety of polymers have been personalized, and their properties fine-tuned to meet specific tissue regeneration needs.^16-18 PLGA polymers are commonly used as regenerative engineering scaffolds, and as materials in FDA-cleared medical devices/implants.^{2, 7} Polyphosphazenes, a unique class of inorganic-organic hybrid polymers, have received a great deal of attention due to their structural diversity, neutral bioactivity, unique erosion, biocompatibility, and miscibility with other clinically relevant polymers.^19-22

In our previous studies, in vitro cytocompatibility and degradation studies were carried out demonstrating the suitability of polyphosphazene-PLGA blends as scaffolding materials for bone tissue regeneration²³. Polyphosphazenes hydrolytically degrade into ammonium phosphates and corresponding side groups (glycylglycine ethyl ester and phenyl phenol) (Figure S1)^23-26. Ammonium phosphate is regarded as an amphoteric compound due to its ability to stabilize the pH of its environment by reacting as a base or as an acid.^{3, 20, 22, 27} When PNGEGPhPh is blended with PLGA, the resulting products have more benefits than the pristine PLGA as the blends generate degradation products with relatively higher pH values than that of PLGA samples²³. This is because the polyphosphazene degradation products can buffer (to some extent) the lactic and glycolic acids from PLGA bulk degradation. Also, the mechanism of degradation of these blends ensues in such a way that the blend samples change from a coherent solid into a porous structure composed of an agglomerate of microspheres. Microsphere formation is attributable to the ability of phenylphenol-containing polyphosphazenes to undergo a hydrophobic aggregation during degradation.²³ This kind of erosion is neither the typical surface nor bulk erosion, and it has been proven to encourage cell infiltration within the interconnected pores and ultimately promote cell growth and enhance cell-material interactions in vitro.²³ Furthermore, since PLGA is already in the market, it is expected that blending it with polyphosphazene-based polymers will facilitate the clearance and commercialization of this class of versatile blend materials.

In an effort to mimic the structure of native bone tissue, this blend system can be fabricated into 3D matrices with similar physicochemical properties as bone tissue²³.

In this study, we evaluated the effects of the PNGEGPhPh-PLGA blends' characteristics on the regeneration of a critical-sized defect model of rabbits and compared the histological response to that of PLGA. Generally, critical-sized bone defects are defects that lack the capability to heal on their own during the patient’s lifetime.²⁸ The size of the bone loss is usually greater than two times the diameter of the long bone, making it extremely difficult for a union to occur despite proper stabilization.²⁹ As previously mentioned, we have illustrated that the polymeric blends between polyphosphazenes and PLGA possessed an erosion profile that enhanced the filtration of cells in vitro.²³ The goal of this study was to understand PNGEGPhPh-PLGA blends degradation patterns and their effects on bone growth in a critical-sized bone defect.

EXPERIMENTAL SECTION

Materials

PNGEGPhPh were synthesized as in our previous study.²⁵ PLGA (50:50, MW = 60800 g/mol) was purchased from Absorbables-Durect corporation and used as received. Chloroform and Tetrahydrofuran (THF) were obtained from Sigma-Aldrich and used without further purification.

Blends of PNGEGPhPh and PLGA ( in 50:50 weight ratio) were fabricated by mutual solvent approach as reported previously in our recent study, and 3D cylindrical matrices were prepared as follows.²³ The blend films were cut into rectangular strips of about 20mm x 7mm x 0.5mm (L X W X T). In a controlled fashion, the films were wrapped up around 1mm PTFE rod to form 3D concentric structures, confined by a PTFE ring with an inner diameter of 10mm. The concentric structures were then immersed in a glass beaker filled with culture medium to drive away entrapped air and to shrink the matrices to the desired size (Figure 1). The resulting 3D scaffolds were 15mm X 5mm in dimension with a central cavity of 1mm (Figure 2). PLGA matrices were prepared by placing polymer in a mold (with cylindrical cavity of 15mm X 5mm) and heating above its Tg.

Figure 1.

Schematic illustration of the design of polyphosphazene-PLGA scaffolds used for in vivo implantation.

Figure 2.

Photo image of cylindrical 3D scaffold fabricated from PNGEGPhPh-PLGA blends.

Methods

Mechanical Evaluation

Tensile tests were carried out on the rectangular polymer strips of 13.2 mm length, 6.6 mm width, and 0.1 mm thickness, using an Instron mechanical testing machine with a crosshead speed of 2 mm/min at ambient temperature and humidity. The testing was performed using five specimens for each sample and in accordance with ASTM D882-12. The ultimate tensile and Young’s modulus were determined from the stress-strain curves.

The materials’ flexural properties were determined using a three-point bend methodology and as stipulated by the ASTM D790-17 standards. Flexural strength (given by the symbol σ) was calculated using:

$$\sigma =\frac{3FL}{2W{t}^2}$$

F means the maximum force applied at fracture, L is the length of the sample (22mm), w is the width of the sample (10mm), and t is the thickness of the sample (2mm).

The scaffolds with 4 mm in length and 3 mm in diameter were used for compressive testing. This testing was accomplished using the Instron mechanical testing machine with a 5 mm/min crosshead speed at ambient temperature and humidity. The compressive properties were then determined from the stress-strain curves.

Rabbit Surgery – In Vivo Study

All surgical procedures were approved by the IACUC facility at the University of Connecticut, CT, USA. Before implantation, the cylindrical scaffolds (PNGEGPhPh-PLGA and PLGA) were exposed to UV light for 15 minutes on all sides and this helps to minimize bacterial contamination. Five male New Zealand White rabbit weighing 3.2-4Kg each were employed for each polymer group and each time point (2, 4, & 6 weeks)(table 1). For the surgical procedure, a lateral incision was made at the forearm of the rabbit and the soft tissues and muscles were dissected to reach the bone. A 15mm critical-sized defect was created at the radius of the rabbit forearm and the polymer matrix was inserted into the space (Figure 3). The muscles and soft tissues were then closed and stitched with sutures to stabilize the implant site (Figure 4a-f). The x-ray images of the rabbits’ forearms were taken before and after the surgical procedures (Figure 5). At designated time points of 2, 4, and 6 weeks, five animals from each polymer group were euthanized. The radial bones of the rabbits were isolated and processed for radiological and histological examinations.

Table 1:

In Vivo Experimental Setup

Total # of Rabbits	30
Time Point (Weeks)	2, 4, 6
Experimental Group	PLGA Control (n=5), PNGEGPhPh-PLGA (n=5)

Figure 3.

Images showing the implantations of PNGEGPhPh-PLGA and PLGA.

Figure 4.

a-c arrows showing the bone-scaffold interfaces of the PNGEGPhPh-PLGA blends and stitched implant area, d-f arrows showing the bone-scaffold interface of PLGA control and the stitched implant region.

Figure 5.

X-ray images of rabbit bone before and after surgical procedure.

Tissue Harvest for Micro-CT and Histology

At predetermined time points of 2, 4, and 6 weeks, rabbits were sacrificed in accordance with established IACUC protocol, and the forelegs were isolated and harvested. X-ray images of the radial bones were generated. The exterior tissues and skins were surgically dissected, and the samples were fixed in 4% paraformaldehyde and dehydrated in 70% alcohol. Some of the fixed samples were scanned using micro-CT (μCT Evaluation Program V6.6) to determine the presence of newly formed tissues. The 3D reconstructed datasets were analyzed to obtain the ratio of bone volume to total volume (BV/TV) within the area of interest. The analysis included new bone formation within the defect area excluding any fragments of native bone. For histological preparation, the paraformaldehyde-fixed tissue samples were cut in the region of bone defect, and samples were processed for embedding in methyl methacrylate (MMA). The embedded samples were stained with Hematoxylin and eosin (H&E), Toluidine blue (TB), Masson trichrome (MT), Goldner’s trichrome (GT) and visualized under a light microscope to evaluate the bone formation and inflammatory response.

Statistical Analysis

All analyses were performed in triplicate or more per sample, and quantitative data were presented as mean ± standard deviation (n≥ 3). Statistical analysis was performed using Microsoft excel. The comparison of the means was performed using a two-sample t test with a significant difference of p < 0.05.

Results and Discussion

Mechanical Properties of Scaffolding Material

The analysis of the physicochemical properties of biomaterials gives insight into the mechanical performance of the newly designed bone scaffolds. Appropriate mechanical performance and dimensional integrity by scaffolding materials play a critical role in determining the overall success of bone tissue regeneration.

Concentrically positioning PNGEGPhPh-PLGA films yielded a mechanically competent scaffold that structurally mimics the natural bone. The samples' mechanical properties and the human cortical and trabecular bones are shown in table 2 and figure 6. For the PNGEGPhPh-PLGA blend, the tensile, flexural, and compressive strengths were 72±1.6, 79±0.96, and 3.52±0.32, whereas the tensile, flexural, and compressive strengths for the PLGA were found to be 35.7±0.5, 58±0.82, and 6.22±0.89, respectively.

Table 2.

Comparison of physicochemical properties of PNGEGPhPh-PLGA, PLGA and native bones.

Material	Tensile Strength (MPa)	Tensile Modulus (GPa)	Flexural Strength (MPa)	Flexural Modulus (GPa)	Compressive Strength (MPa)	Compressive Modulus (GPa)	Tg (°C)
Blend	72±1.6	2.85±0.35	79±0.96	7.8±0.38	3.52±0.32	0.25±0.11	70
PLGA	35.7±0.5	1.48±0.6	58±0.82	2.96±0.5	6.22±0.89	0.39±0.21	40
Cortical	52 – 150	7 – 30	50 – 150	-	67 – 230	17 – 20	-
Trabecular	1 – 10	0.05 – 0.1	10 – 20	-	0.2 – 12	0.02 – 0.9	-

Figure 6.

Graphical representation of the mechanical properties of PNGEGPhPh-PLGA blend and PLGA. The blend showed superior tensile and flexural strengths, while its compressive properties were comparable to that of the pristine PLGA. Data are expressed as means ±SD (n=5). p < 0.05.

The PNGEGPhPh-PLGA materials exhibited excellent tensile and flexural properties as these properties were superior to that of the pristine PLGA and similar to the native bone. Besides, the blend's compressive strength was comparable to that of the PLGA and within the range of the trabecular bone but fell short of that of the cortical bone.^{1, 30} Similar trends were also observed in the results for the Young’s, flexural, and compressive moduli.

Interestingly, the mechanical evaluation results suggest that the newly designed scaffolds were mechanically suitable for bone regenerative engineering because their initial tensile strength matches that of the cortical bone. It is worth mentioning that the mechanical properties of degradable biomaterials progressively drop as they undergo degradation. Hence, adequate initial mechanical properties are ideally and optimally necessary.

General Observation of the Implantation

Visual observation of the implant regions at 2-, 4-, and 6-week time points after surgical implantations indicated no gross tissue disturbance, swelling or physical impairment for the PNGEGPhPh-PLGA system (Figure 7). However, two of the fifteen rabbits implanted with PLGA showed a mild inflammatory response in the early days of implantation consisting of swelling, which normalized after a few days. Also, there were no gross signs of systemic or neurological toxicity for the rabbits with both PNGEGPhPh-PLGA and PLGA implants.

Figure 7:

Harvested PNGEGPhPh-PLGA-implanted bone showing intact matrix after 4 weeks.

X-ray, Micro-CT, and Histology

The x-ray analyses were used to examine new bone formation around the defect region as the x-ray images of the rabbits’ forelegs were obtained at each time point and for each matrix. The images showed a progressive increase in radiodensity (or radiopacity) around the defect areas, thought to be consistent with new bone formation and growth.³¹ The earliest time point (second week) results showed increased radiodensity around the edges of the pre-created defects. The process continued with more areas of the defects, demonstrating increased bone formation after four weeks, and this is indicated on the X-ray images in figure 8. At the 6-week time point, the PNGEGPhPh showed high levels of radiodensity at the rabbits’ forelegs compared to the PLGA. This is consistent with the results of our previous in vitro studies as the relatively mild and less acidic degradation products (and the morphological changes) of PNGEGPhPh-PLGA might have allowed more enabling conditions and environment for new bone formation and its integration with the native bone tissues (Figures S2 & S3).

Figure 8:

X-ray images of the rabbit forelegs before or after implantations for PNGEGPhPh-PLGA and PLGA at different time points. X-ray results revealed increased radiodensity at the defect site consistent with the promotion of new bone formation.

The reconstructed images of the micro-CT were cut into rectangular shapes that featured the pre-created defect region, and the newly formed bones were monitored as it filled and covered the defect gap with time. The micro-CT images of the fourth and sixth weeks showed that newly formed bone for PLGA only occurred at the edges of the defects while that of the Polyphosphazene-PLGA blend was not only prominent at the edges but bone formation was seen simultaneously at the centers of the defects (Figure 9). Consequently, there were more areas covered by newly formed bone via the PNGEGPhPh-PLGA scaffolds than that of the PLGA scaffolds. Bone volume fractions (BV/TV) were analyzed, and the volume of newly formed bone for PNGEGPhPh-PLGA was found to be significantly higher than that of the pristine PLGA at all time points (Figure 10). There was a linear relationship between the bone volume fraction and the time points as the bone volume, and volume fraction increased with increasing number of weeks. The PNGEGPhPh-PLGA implants witnessed a faster formation of new bone volume after 4 weeks in comparison to PLGA. This particular analysis suggests that there was relatively faster bone regeneration within the polyphosphazene-based matrix than the PLGA. The rapid regeneration exhibited by PNGEGPhPh-PLGA could be attributed to the blends’ unique features such as self-neutralizing effects (Figure S2), peptide release (which may act as nutrients for local cells) (Figure S1)^23-24, and microsphere-forming tendencies upon degradation (Figure 11). The peptide ester cosubstituent of the PNGEGPhPh is released during its hydrolytic breakdown.

Figure 9:

MicroCT images of rabbit radial bone defects repaired by PNGEGPhPh-PLGA and PLGA scaffolds. PNGEGPhPh-PLGA implants visually demonstrated higher levels of bone formation.

Figure 10:

Graphical representation of BV (bone volume)/TV (total volume) of PNGEGPhPh-PLGA and PLGA scaffolds. The volume fractions of the newly formed bone increased with time for both implants but the rate of bone formation was faster for polyphosphazene scaffolds. Data are expressed as means ±SD (n=5). p < 0.05.

Figure 11:

SEM images of PNGEGPhPh-PLGA and pristine PLGA showing different morphologies after degradation. PLGA polymers typically undergo buck erosion, where degradation occurs throughout the whole material equally (from inside to outside) and with a total breakdown of polymer chains and molecular weight. The blend’s degradation profile led to the formation of microspheres, which aggregated into interconnected 3D porosity. The unique degradation characteristics may have contributed to the faster rate of bone tissue regeneration for PNGEGPhPh-PLGA. Scale bar = 10μm.

H&E and Toluidine blue staining were performed on the regions of interest to analyze the potential histological changes of the newly formed bone tissues within the regenerated bone defect areas. Masson’s trichrome and Goldner’s trichrome stains were further employed to complement and confirm the H&E and TB analyses. As shown in figures 12a & 13a the MMA-embedded tissue sections were stained with H&E, and interestingly, all the scaffolds of the PNGEGPhPh-PLGA and PLGA groups showed no inflammatory responses. Additionally, substantial new bone was observed in the PNGEGPhPh-PLGA group; however, new bone was found to occupy a small area within the region for the PLGA group. There was no apparent and significant new bone observed in the PLGA group at the 2-week time point.

Figure 12:

Comparison of (a) H&E (b) Toluidine blue (c) Masson trichrome and (d) Goldner’s trichrome stained images of polyphosphazene-PLGA blend matrices for 6-week critical-sized defect model. The blend matrices showed more promising translational capability as they exhibited more rapid bone regeneration than the PLGA scaffolds 6 weeks post-surgery and implantation. NB indicates new bone, HB indicates host bone, and RS indicates residual scaffolds. Scale bar = 2mm(first column) and scale bar = 1mm (second and third column).

Figure 13:

Comparison of (a) H&E (b) toluidine blue (c) Masson trichrome and (d) Goldner’s Trichrome stained images of PLGA samples (6 weeks after surgery). Less bone formation was observed for PLGA group as compared to the blend group 6 weeks post-surgery and implantation. NB indicates new bone, HB indicates host bone, and RS indicates residual scaffolds. Scale bar = 2mm(first column) and scale bar = 1mm (second and third column).

Figure 12b & 13b shows the Toluidine blue-stained images of samples of the two different matrix groups. The Toluidine blue stain is mainly used to identify connective tissues, precisely collagen type 1, which is abundant in bone. The light-colored stains represent the intercellular collagen fibers, which filled up the residual scaffolds, while the newly formed bone was dark-colored stains, which is due to the abundant presence of compact collagen. Similar to H&E staining results, the new bone regenerated within the PNGEGPhPh-PLGA matrix existed both in the edges and in the middle of the defect region, whereas there was a minor new bone formation in the PLGA, which only occurred around the edges of the defect. These results were in agreement with the outcomes of the Masson trichrome and Goldner’s trichrome analyses (Figures 12c, 12d, 13c, & 13d) as there was a sharp contrast between the host and new bone tissues. The deep blue in MT and deep green in GT represent the mature bone matrix, while the light-colored blue and green stains correspond to the immature new bone matrix for MT and GT respectively.

Besides osteoblast, there existed osteoclast cells within the region of the newly formed bones, and the presence of these two cell phenotypes is an indication of active bone remodeling. The radiological and histological results demonstrated that the polyphosphazene-containing 3D scaffolds promoted rapid bone tissue regeneration in vivo, and as such, they are promising candidate materials for regenerative engineering applications.

CONCLUSION

In this study, we found that 3D matrices composed of polyphosphazene-PLGA blends exhibited excellent physicochemical and in vivo performance as the matrix promoted the formation of new bone in a rabbit critical-sized radial defect model. Based on radiological and histological analyses, there was accelerated bone regeneration and a mild inflammatory response associated with the PNGEGPhPh-PLGA implants, as compared to PLGA. Overall, this study demonstrates that blends of peptide-substituted polyphosphazene and PLGA are biocompatible with bone tissue and may prove to be a viable biomaterial for matrix-based bone regenerative engineering. The difference in the erosion mechanism of PNGEGPhPh-PLGA and PLGA may explain the difference in their tissue responses. This suggests that the intrinsic microsphere-forming properties of the PNGEGPhPh-PLGA system may be an essential and advantageous factor for adapting this class of biomaterials to specific musculoskeletal tissue regeneration needs.

• Degradation mechanisms of peptide-containing polyphosphazene

• Self-buffering effect of polyphosphazene-based degradation products

• Cell viability of polyphosphazene-based biomaterials