Effects of hydroxyapatite on PMMA-HAp cement for biomedical applications

1. Introduction

Polymethyl methacrylate (PMMA) is a useful material for minimally invasive spinal percutaneous procedures applied to vertebroplasty and kyphoplasty. These procedures involve the percutaneous injection of PMMA’s bone cement into a collapsed vertebral body. The cement in the vertebroplasty procedure bonds fractured bone pieces stabilizing the vertebral body. Also, the cement in kyphoplasty has a similar function; however, a balloon is inflated previously to restore the vertebra anatomy [1]. In general, the percutaneous techniques have been used for structural recovery in osteoporosis bone disease [2].

Pathologic fractures often occur under normal physiologic stress in patients with spinal metastasis. The fracture can occur and entail to a partial or total damage of the anterior vertebral body [3]. In oncology cases, bone metastasis can significantly affect a patient’s quality of life due to the debilitating pain, fractures or even paralysis by spinal cord compression [4]. Radiovertebroplasty has been early suggested following a similar procedure of vertebroplasty; however, with the insertion of a radioactive cement in situ [5]. Radiovertebroplasty has not yet fulfilled the conditions of clinical application. It has already been studied by means of a computational simulation in which the spatial dose delivered from the radioactive bone cement with Sm-153 or Ho-166 targeting hydroxiapatite (HAp) was addressed in a vertebral phantom [6–8]. The absorbed dose was also evaluated by Monte Carlo Computer Code (MCNP5) with P-32, Ho-166, Y-90, F-18, I-125 and Tc-99m isotopes homogeneously distributed on the vertebrae. Both studies showed that the spreading of the spatial absorbed dose distribution was limited to the range of the beta emitters, which favors the radiotherapy of metastatic lesions [6–10].

The synthesis and characterization of calcium phosphate loaded with Ho-166 and Sm-153 were also investigated. The nuclear characterization of HAp doped with Sm-152 and Ho-165, their decayment processes, radionuclide contaminants and the activation process in a research reactor have been addressed, which demonstrated the viability of producing radioactive cement in a low-flux reactor-type. Although the nuclear proprieties of HAp-¹⁵³Sm and HAp-¹⁶⁶Ho were presented [6], thermal and mechanical properties of the composite made of PMMA-HAp has not yet been addressed.

There are other materials used as bone cements that are worth mentioning. An injectable gel from squid pen Chitosan has been studied for bone tissue engineering applications [11]. Chitin or Chitosan is a colloid with three different structures. One is very stiff with structures in a sheet shape in which each sheet is linked by hydrogen bonds while others have a low packing crystal structure and weaker hydrogen bonds. The mixture of calcium phosphate compounds (CPC) with Chitosan solution has already been proposed in appropriate ratios. Such a biocomposite not only supports and increases the bone regeneration ability of the gel system, but also increases the mechanical properties of the gel system [12]. The gelation time was less than 3 min at 37 °C without insertion of CPC, while the gelation time decreased with the increase of the concentration CPC. The sol-gel transition in Chitosan/Group of phosphates (GP) solutions is a physical process governed by different interactions such as hydrogen bonding, hydrophobic effect, electrostatic repulsion and ionic cross-linking [13]. The mechanical properties of the material improves if the Chitosan concentration is increased; however, it would likely impair the clinical injecting procedure. The study of the mechanical characteristic of calcium phosphate compound mixtures with Chitosan, especially HAp, is of great interest, particularly to compare it with bone cement PMMA-HAp, since in principle both were designed for the same purpose [11]. In the case of PMMA, there is no chemical reversion after the polymerization process (PP), although PP is temperature dependent and cannot occur below 8 °C.

Thermal, morphological and mechanical properties of PMMA electrospun nanofibers have already been studied [14,15]. Differential scanning calorimetry (DSC) was carried out to determine the thermochemical behavior and phase change by increasing the PMMA/NaCl solution concentration. The results showed that the T _m (melting temperature) for a first glass transition was almost 116.19 °C to two samples with 12.5 wt% and 15 wt% concentrations, i.e. invariant with NaCl concentration. Nonetheless, the second glass transition temperatures were 121.58 °C to 12.5 wt% and 123.06 °C to 15 wt% for a second thermal anomalous, confirming that the crystallinity of the fibers of PMMA/NaCl increases with concentration [14,15]. On the one hand, a higher value of the glass transition temperature T _g yields stable thermally fibers [14,16]. An initial decomposition temperature of the PMMA powder at 140 °C and to PMMA/NaCl fibers 302 and 320 °C, respectively, was also reported. This phenomenon is caused by the sub-products during the thermal degradation process [14,17]. Other studies on thermogravimetry (TG)-differential thermal analysis (DTA) were developed in HAp to examine the decomposition process attributed to nitrates and urea, thus this being the largest weight loss. This phenomenon occurs near 200 °C. It is important to note that the methods of synthesis of HAp in both works are different. On the other hand, many positive aspects achieved when heat treatment at 800 °C was included [18]. Another study presents an exothermic peak due to the combustion of organic components around 300 °C, provided by TG-DTA  [19].

The mechanical property of nanofibers in ultimate tensile strength (UTS) was about 1–3 MPa. The modulus of the parallel layers in the nanofibers was 52.3 ± 5.2 MPa. Also, when the layers were arranged in cross, the modulus was 26.1 ± 4.0 MPa [14,20]. Based on previous work, Akhtar et al. suggested that the nanofibers PMMA/NaCl, with an optimal mechanical behavior, can be used for applications in in vitro cell growth. The cells should be deposited on these PMMA nanofiber’s looms for biomedical and pharmaceutical areas [14]. Studies on the morphology and strengths of composite bone cement showed a compression of 71 MPa with a large amount of pores with a diameter length over 100 μm [21].

Concepts of biomechanics associated with bone properties have previously been reviewed [22], in which the biomechanical characteristics of the porous of the bone tissues could be appreciated [23]. In this review, the stress vs deformation profiles in an elastic and a plastic regions could be classified. It should be noted that there are many rheological and mechanical characteristics to be studied in porous composites; but, for practical purposes, it remains useful to study the properties of resistance and Young’s modulus of the bone or bone substitute materials. In natural tissue, resistance values for cortical bone range from 167–213 MPa and Young’s modulus range from 14.7–34.3 GPa [23]. For spongy bone, the resistance varies from 1.5–9.3 MPa and the Young’s modulus varies between 10–1058 MPa [23]. Such data can be addressed to qualify synthetic biomaterials as bone substitutes in comparison to natural bone.

For the tissue engineering proposal, the development of a composite mimicking trabecular bone continues to be a challenge. We aimed to investigate the PMMA-HAp system that approaches the trabecular bone characteristics in its end into the vertebrae. To do this, the main goal of the present paper was to study the thermal and mechanical properties of bone cements made of PMMA-HAp composite in the cold condition (non radioactive) in function of the HAp concentration and to study the effects in the porous production.

2. Materials and methods

3. Results

3.1. Differential scanning calorimetry (DSC) profiles

The DSC thermal profiles for the polymethyl methacrylate (PMMA) (copolymer)-HAp and a PMMA-HAp system in the first scanning calorimeter is depicted in Fig. 1. Two anomalies were observed. The anomalies in DSC profiles are in a temperature domain between 100 to 120 °C and 130 to 160 °C, respectively. Figure 1a depicts the anomalies for the first four concentrations which held a higher proportion of PMMA, which weaken with the increase in the ratio of HAp in the sample. The first anomaly is of no interest in this study because the occurrence of water loss or by-product formation are unrelated to glass transition in the analysis. The second anomaly was observed at the first four concentrations (x ₁, x ₂, x ₃ and x ₄). The dotted lines guide the reader to the location of the events described. The phenomena that occurred in the samples was exposed to the heat treatment in which the baseline was corrected for each concentration. Another relevant feature, in the 120 to 130 °C domain, is a possible T _g glass transition reported previously [14,15]. Here, the crystallization phenomena were understood as the ratio between PMMA/HAp. Indeed, the biophosphanated calcium matrix should increment the crystallization level in the bone cement. The following peaks from the first four concentrations correspond to the curing process before the melting of the cement in the temperature domain close to 140 °C. For the polymerized system in Fig. 1b, it was difficult to identify the glass transition temperature, probably because of the different overlapping processes that occur in this temperature domain.

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Fig. 1.

DSC heating curves for the six [x] concentrations of the (1-x)PMMA-xHAp binary system in the interest temperature ranging between 60–180 °C approximately. The heat flux is presented in [mJ/s]. All curves were normalized with the maximum heat flux [𝛷_max in mJ/s], i.e. 𝛷_max = 0.10078 mJ/s and 𝛷_max = 0.21781 mJ/s respectively. (a) PMMA (copolymer)-HAp system with glass transition at T _g close to 125 °C on average; and (b) PMMA-HAp polymerized system.

3.2. Thermogravimetric analysis (TGA) profiles

The TGA thermal profiles for the PMMA (copolymer)-HAp and PMMA-HAp systems are depicted in Fig. 2. The TGAs showed a significant weight loss after 100 °C for the x _i (i = 1,6) concentrations that are associated with water loss. The maximum weight losses of 4 to 7 % occurred after T _g glass transition.

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Fig. 2.

TGA heating profiles for x _i (i = 1,6) concentrations of the (1-x)PMMA-xHAp binary system in the interest temperature ranging between 45–180 °C approximately. (a) PMMA (Copolymer)-HAp, and (b) PMMA-HAp polymerized system.

There is no weight loss equivalent to PMMA (Copolymer)-HAp (Fig. 2a) on the PMMA-HAp polymerized system (Fig. 2b), which presents less than 1%. A possible explanation might be a higher retention of water and the presence of other components that have no participation in the polymerization process or a compound that was used as a catalyst. In Fig. 2b, a weight loss close to 4% to 7% occurred because of the formation of internal pores in the system. Also, significant weight losses were observed in Figs 2a and 2b at 120 to 160 °C intervals. Above this temperature, polymer melts.

3.3. Mechanical resistance

Two assays were performed. The first one addressed the mechanical behavior of the cement as a function of the concentration of HAp. The second one addressed the dependence on the concentration simultaneously with the test block thermally treated around the glass transition temperature.

Figures 3a and 3b depict the profiles of the resistance to compression of the composites as a function of their x, x _i (i = 1,5) concentrations in the PMMA-HAp system. A geometric factor was considered for each block in the group with the same HAp concentration [x _n ]. The analysis of the compressive strength was obtained using the slope in relation to the baseline technique. The graphs in Fig. 3 show non-conforming cases of samples from a strain 0.3 mm/mm in the plastic region, setting new dimensions for each block.

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Fig. 3.

Compression profiles for the mechanical assay on the blocks with x _i (i = 1,5) concentrations of the (1-x)PMMA-xHAp binary system. (a) PMMA-HAp polymerized system without thermal treatment with a fracture domain at 0.30–0,42% compressive strength; and (b) PMMA-HAp polymerized system with thermal treatment of 30 min at T _g glass transition temperature.

Table1 summarizes the data of strength in function of the temperature in the thermal treatment and concentration of HAp. It is observed that the glass transition temperature and weight losses were relatively stable for the concentration parameter x of 0.0 up to 0.5, thus we were confident that the chemical and mechanical processes that were developed in each sample with distinct concentrations would be equivalent.

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Table 1

Thermomechanical properties of the (1-x)PMMA-xHAp system

x n	Concentration [x]	Glass transition T g [°C]	TGA weight loss [%]	Compressive strength [MPa]	Compressive strength with thermal treatment [MPa]
x 1	0.00000	127.15	4.02	64.44	84.00
x 2	0.02167	126.93	6.80	60.45	82.77
x 3	0.09062	124.67	4.33	87.60	70.13
x 4	0.16619	124.35	4.89	41.40	70.07
x 5	0.50000	126.27	<1	19.50	3.71–7.37
x 6	1.00000	—	<1	—	—
∗	Cortical or compact bone ∗ and
	∗∗cited by Caeiro et al. [23]			167–213	—
∗∗	Trabeculae or spongy bone			1.5–9.3	—

3.4. Stereo surface images of PMMA-HAp samples

Figure 4 shows the surface of the biophosphate-free block without thermal treatment (x ₁) (Fig. 4a) and a block with the highest concentration of HAp (x ₅) with thermal treatment to the T _g temperature.

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Fig. 4.

Stereoscopic images of the block surfaces. (a) PMMA only or x ₁ concentration without thermal treatment; and (b) x ₅ concentration parameter with thermal treatment for 30 min at the T _g glass transition temperature.

The PMMA-HAp binary system in the study formed the basis of various commercial bone cements. The polymethyl methacrylate (PMMA) provided two liquid and powder presentations whose moisture yields the polymerization process aggregating the monomers and PMMA microparticles. At the PMMA- HAp composite preparation, the water solution modulated the phases of the acrylic resulting in an increase in the polymerization time (PT). The instrumentation and material cooling procedure also increased the PT due to the reduction of the intensity of the polymerizing effect that is proportional to the ambient temperature.

Our research team is committed to design a system that potentates mineralization processes in situ, especially those associated with spongy bone tissue. The production of pores in situ after implantation is a difficult task since natural scale pore generations have a high degree of randomness. The pore generation was understood as a spontaneous defect in the material. The pores should be connected or have thin walls in which fluid can be diffused, allowing plasma and bone narrow diffusion, in which angiogenesis and other vascularization processes can be performed facilitating the natural biological processes.

Likewise, in differential scanning calorimetry (DSC) analysis (Fig. 1b), the stability of the baseline was corrected to improve the understanding of the thermal phenomena that were developed in the material. It is important to emphasize that all thermal profiles were normalized with the maximum thermal power to evaluate how the present anomalies behaved as a function of the concentration of HAp in the cement. Another important aspect of the DSC analysis of our systems was to confirm the existence of T _g glass transition in the temperature interval that was already mentioned by other researchers [14], which question how the mechanical properties of the composite are when they have been treated thermally.

The second-order transitions in thermogravimetric analysis (TGA) (Fig. 2b) suggested the possibility of increasing the polymorphic characteristics of the system and, in turn, promoting more defects associated with its porosity. Some reactions with gas-releasing may start after the glass transition where the polymer chains interact with (OH)⁻, with groups producing water vapor and CO₂. These low weight losses are depicted in Fig. 2a and confirm that calcium matrices are only in the heterogeneous phase in the system which reduces the gas-releasing process due to the interaction of the (OH)⁻ groups and the polymer. The polymerized blend (Fig. 2b) remains a much more intricate heterogeneous composite than calcium matrices and polymer chains.

Considering the mechanical resistance depicted in Fig. 3a, the region between the dotted lines shows a tentative of defining the strain domain where the fracture of the material occurred. In principle, the modulus of compressive strength decreases with an increasing concentration of HAp. This reasonable argument allows us to understand that the thermoplastic properties of the material are reduced and the plastic region is considerably smaller between the concentrations x ₁ and x ₅. In Fig. 3b, one can observe that the plastic region is smaller for the concentrations x ₄ and x ₅, whereby it becomes clear that the material is more fragile. It is important to recall that the blocks were heated to the glass transition temperature in order to examine the mechanical feature of the material after glass-transition temperature treatment. The weight losses were estimated from the glass-transition temperature to a post-melting polymer temperature where the TGA profile is stable. The compressive strength (CS) for the block of x ₁ concentration parameter was one-third of the CS mean values to human cortical bone; while it was 15-fold for x ₁ and 4-fold for x ₅ higher to mean value to the human spongy bone. For the blocks with thermal treatment, the CS presents close to 30% higher than that without thermal treatment; although for the concentration x ₅ the CS values is in the range of the spongy bone tissue.

In the compression resistance analysis, the CS values were slightly lower than the actual compression values. The elasticity or Young’s modulus calculated as a function of the concentration parameter varied from 1.87 GPa for PMMA-HAp x ₁ to 0.21 GPa for x ₅ decreasing steadily as the proportion of HAp [x] increases in the system. An elastic behavior for PMMA- HAp was observed diverging from the cortical bone tissue propriety and closer to the spongy bone tissue, as expected [23]. The toughness can be calculated as the area under the curve of the stress vs strain plots. The average of the compressive strengths for each concentration showed values ranging from 64.54 to 19.50 MPa for x ₁ to x ₅ concentrations. The CS values for the thermal pre-treated blocks were of 7.37–3.71 MPa. There were two triplicates set for x ₅ modulating with and without water. The CS average was 3.71 MPa for x ₅ modulated with water and 7.37 MPa without modulation. It is observed in Fig. 4b that the outset of the defect formation could fit the intention of the in situ pore generation.

Subsequent studies will enclose both transport phenomena and rheological analyses. Those investigations can predict the ability of plasma diffusion in PMMA-HAp. The knowledge of the viscosity variability of the PMMA-HAp during the curing process in function of the polymerization time and water content is relevant to find a better way to implant cement. An important prospect for our study will be to focus on the PMMA-HAp biocompatibility degree, taking into account that the PMMA polymer has a relative biocompatible for biomedical applications. On the other hand, we know that HAp together with collagen are natural components present in bone tissues that are produced and reabsorbed consistently within the conceptual framework of permanent restoration of the skeletal system.

5. Conclusions

In this study, it was possible to establish a direct relationship between the thermoplastic properties of the PMMA+HAp acrylic system and its mechanical properties so that changes in function of the biophosphate concentrations could be depicted. The concentrations of the HAp constitutes could be used to optimize its mechanical qualities for the purpose of body metastasis application. The biomechanical results for the concentration x ₅ presented compressive strength values between 3.71–7.37 MPa, holding similarity to the spongy bone presented by Caeiro et al. with values of 1.5–9.3 MPa, according to Table1 [23]. The greatest innovation reached in this work was the improvement in bone cement syntheses matching the desired biochemical properties represented by the increase of the porosity of the material, such that the bone marrow physiological diffusion may occur. An increase in the HAp amorphous polycrystalline structure favors the segregation of the different segments of the polymethyl methacrylate (PMMA) polymer due to its hydrophobic nature similar to the reactions among amphiphilic biomolecules and water. The pore sizes are perceptible to the human eye and are very close to those found in the trabecular bone tissue itself. This phenomenon was optimized with the thermal treatment carried out on the samples in the glass transition temperature approaching 127 °C. On the other hand, the possible dehydration reactions that occur in the material, enhance the formation of more pores and increase their size, possibly because of the subproduct releases, water and CO₂ in the heterogeneous mixture.