Introduction

A biomaterial used as a medical device must have intrinsic characteristics that allow its application with a long-term duration and without rejection by the human body [1, 2]. Among the polymeric biomaterials that have been investigated for biomedical devices is poly (lactic acid) (PLA) [3]. PLA is an aliphatic bio-based polymer, preferable in biomedical applications due to its biocompatibility, biodegradability, and low toxicity [4]. PLA is used in a wide range of biomedical applications, such as stents [5,6,7], surgical sutures [8, 9], and screws [10], among others. However, the general criterion for selecting a polymer for its use in a medical application is to balance the mechanical properties and the required degradation time. Therefore, the selection requires an in-depth characterization of the changes in strength and stiffness of the polymeric material during degradation [7], to maintain mechanical stability during this stage. Understanding the PLA properties is the key to achieving suitable chemical and biological properties.

The degradation behavior of PLA is a vital feature and a significant reason for its use in medical applications [7]. However, it presents some limitations, such as hydrophobicity and inadequate mechanical properties for bone repair [11, 12]. Amorphous PLA polymer has low tensile strength, higher elongation at break, and more rapid degradation time than semicrystalline PLA, making it more attractive as a drug delivery system [7, 13]. Specifically, in orthopedic and fixation devices, an advantage of a PLA medical device is that it does not have to be removed, avoiding a second surgery and reducing costs because it degrades simply by hydrolysis without using enzymes or catalysts [7, 14,15,16].

Although the biodegradability of materials used in the manufacture of medical devices is desirable, the polymer degradation rate must be customized to be safe without cytotoxic or inflammatory reactions [17,18,19]. Velioglu et al. [20] evaluated the potential of utilizing PLA in medical devices. They studied the attachment and proliferation of human bone marrow mesenchymal stem cells of 3D-printed poly(lactic acid) scaffolds concluding that PLA scaffolds appear as an alternative for bone repair and regeneration. The key to medical device success is to balance the presence of additives, the geometry of the device, and its location to adapt an implant for slow degradation and transfer stress to the surrounding tissue as it heals at the appropriate rate. Another requirement for materials intended for contact with blood, e.g., medical devices, is to evaluate their hemolytic properties. The presence of hemolytic material may produce increased levels of free plasma hemoglobin capable of inducing toxic effects or other consequences which may stress the kidneys or other organs [21].

On the other hand, some authors have studied the hydrolytic degradation of PLA in PBS solution at 37 °C [22, 23] and under different temperatures. For example, Gorrasi and Pantani [24] reported the hydrolysis process during 60 days of PLA with different D-isomer content by immersion in distilled water at 58 °C and a pH of 5.7. The degradation medium was replaced every day. Around day 30, they found that PLA 4060D (PLA with the highest D enantiomer content) showed the fastest rate of weight loss with a value of 70%. Likewise, Wei et al. [25] investigated the hydrolytic degradation behavior of PLLA grafted with poly(β-myrcene) (PM) rubber during 90 days, registering a mass loss of the PLLA homopolymer around 73.9%, while the PM-g-PLLA copolymers registered less than 5%.

Nevertheless, PLA exhibits fragile behavior and low elongation [26,27,28]. The polymer blend is an efficient and cost-effective method for tailoring and modifying properties [29]. In particular, PLA generates a new degradation profile that depends on the type of polymer added, the relationship, and the morphology of the blend [8]. In addition, PLA can be blended with other polymeric components to cover most required properties in worldwide application requirements [17]. Furthermore, two significant challenges to stabilizing PLA are its water permeability and autocatalytic reaction [17]. Therefore, melt blending is ideal for modifying polymer properties and does not employ any solvent-avoiding non-desirable residual solvent toxic effects, which medical devices require [30]. In addition, minimal attention has been given to studying blend stability under different hydrolytic degradation conditions [7].

Moreover, Liu et al. [31] investigated the in vitro degradation behavior of PLLA/aragonite and PLLA/vaterite blends. They observed that blending the PLLA can decrease its rate of hydrolysis. Navarro et al. [32] studied in vitro biodegradation of PLA/PCL blends at 37 °C for 140 days in a PBS solution replaced once a week. They found that under the conditions established in vitro test, PLA degrades by chain scission of polymer chains with a decrease of molecular weight but without weight loss. Wongwiwattana and Thomas [33] conducted degradation tests in PBS solution for 260 days at 37 °C of PLA/PCL, indicating that PLA undergoes bulk degradation.

The obtention of PLA/EVA and PLA/SMMA binary blends has been previously investigated separately [30, 34,35,36,37,38,39]. In particular, PLA/EVA blends have already been considered for medical applications [30, 37, 40]. There are few reports about PLA/PMMA blends as an option for medical applications [41].

In the previous study, we investigated the effect of mixing PLA with EVA and SMMA (EPS) obtained by melt blending to improve its mechanical behavior regarding elongation at break and stiffness [34]. The addition of EVA improved the elongation at break of PLA, while SMMA contributed to preserving the stiffness. In addition, the effect of processing and incorporation times of the components in the mechanical properties was studied [34]. In contrast, in this research, we explore these polymer blends with potential applications in medical devices. This study aimed to determine the mechanical properties and morphology of the EPS blends during hydrolytic degradation, specifically in phosphate buffer solution (pH 7.4) at 37 °C. Furthermore, the cytotoxicity and hemolytic effect of these blends were evaluated for their application in medical devices.

Materials and Methods

This research work was authorized by the Ethics Committee of the Universidad Autónoma de Nuevo León, School of Medicine with registration number BI19-00003 and Dr. José Eleuterio González University Hospital of the Universidad Autónoma de Nuevo León. All animals and human donors were treated according to ethical standards.

Materials

Poly (l,d-lactide) PLA Ingeo 4060D, (8%–10%) D [42], from NatureWorks LLC, USA, with an average molecular weight Mw—190 kg/mol, ρ = 1.24 g/cm3 and glass transition temperature Tg = (55–60) °C.

Poly(ethylene-co-vinyl acetate) (EVA), commercial name EVA 2810-A ATEVA CELANESE with 28% vinyl acetate, ρ = 0.949 g/cm3, Tm = 73 °C, MFI = 6. Poly(styrene-co-methyl methacrylate) Poly(S-co-MMA), SMMA NAS®30, ρ = 1.090 g/cm3, Tg = 103 °C, MFI = 2.2, from Ingeos Styrolution Group GmbH, Germany with the content of styrene (70–90%) and methyl methacrylate (10–30%). Phosphate salt (PBS) capsules from Sigma-Aldrich, bone marrow mesenchymal stem cells (BM-MSC) were obtained from BALB/c mice. CO2 gas, DMEM/nutrient mixture, FBS, gentamicin, and amphotericin (Gibco; Thermo Fischer Scientific, Waltham). Cell culture flasks (Corning, Inc.). MMT Solution (Cell proliferation Kit I, Roche, Basilea Switzerland). Isopropanol, methanol, HCl (Sigma Aldrich), and BD Vacutainer® heparin tubes (Becton Dickinson).

Melt Blending

The ternary polymer blends EVA/PLA/SMMA (EPS) were prepared following the procedure reported in our previous paper [34]. The formulations chosen for the present study are summarized in Table 1. The ternary polymer blends EVA/PLA/SMMA (EPS) were prepared in a Brabender internal mixer (BB) [DDRV501, C.W. Brabender Instruments Inc., Hackensack, NJ, USA] at a temperature of 180 °C, see Fig. 1a. Table 1 shows the formulations of manufactured polymer mixtures. To remove moisture that polymers might contain and prevent possible hydrolysis degradation during processing, the pellets were dried previously at 60 °C for 12 h employing an Isotemp model 281 vacuum oven.

Table 1 Formulations of polymer blends produced
Full size table
Fig. 1
figure 1

Experimental design. a Blends preparation, b obtention of probes and films, c array of probes for hydrolytic degradation at 37 °C in PBS, and d characterization at different degradation times

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Films and Probes Elaboration

The blends obtained were cooled to room temperature and ground using a blade mill, Fritsch model Pulverisette. To determine the mechanical properties and to evaluate the cytotoxicity of the blends, thin films (thickness of 0.15 mm) and type V probes (ASTM D638) [43] were compression molded using a hydraulic compression molding press (Carver Inc. Wabash, IN, USA) at a temperature of 180 °C, the applied force was 5 metric Ton, see Fig. 1b.

Hydrolytic Degradation

The experiments were conducted in phosphate-buffered saline solution (PBS) (pH = 7.4 ± 0.1, Fisher Scientific) at a temperature of 37.0 ± 0.1 °C, see Fig. 1c. Twenty-one probes (ASTM D638 Type V) of each formulation (EPS1, EPS2, EPS3, PLA5, and PLA15) were placed vertically inside acrylic boxes, and a PBS solution (2 L) was added, covering 1 cm above the sample. In addition, three specimens of each formulation were removed from the boxes at 1, 7, 21, 35, 42, 84, and 168 days for evaluation. The probes were incubated in a forced air oven at 37.0 °C until 168 days. The PBS solution was not changed during the experiment, and a box was utilized for each sample, see Fig. 1c.

Obtention of an MSC Primary Culture

Bone marrow mesenchymal stem cells (BM-MSC) were obtained from BALB/c mice. In addition, three mouses were placed inside a CO2 chamber (70% vol/min) to sacrifice them. Immediately after, their femurs and tibias were dissected under aseptic conditions. Epiphysis was removed, so BM was obtained using a hypodermic needle (27 gauge) by perfusing the shafts inner barrel with 500–1000 μL DMEM/Nutrient Mixture F-12 (DMEM F-12) medium supplemented with 10% FBS, 100 μg gentamicin/mL, and 2.5 μg amphotericin B/mL (Gibco; Thermo Fisher Scientific, Waltham, MA, USA), and poured into T-25 cell culture flasks (Corning, Inc.). Cells were incubated for 24 h at 37 °C in a 5% CO2 humid atmosphere. The culture flasks with the adherent cells were replenished with 4 mL of fresh supplemented DMEM F-12 medium and incubated before the cell monolayer reached 80% confluence (7–10 days). The expended medium and the non-adherent cells were discarded.

Characterization

Water Absorption and Remaining Mass

The initial weight (Wi) of the samples was registered prior to hydrolytic degradation using a 4-digit analytical balance (OHAUS Adventurer AR3130). Subsequently, the weight of three samples at a specific time (1, 7, 21, 49, 84, and 168 days) was registered. Then, samples were weighed to determine their wet weight (Ww) by removing excess water with Kimwipes (Kimtech science, Kimberly-Clark). Later, these probes were rinsed with distilled water and dried at 25 °C for 72 h at a set vacuum of 0.8 bar until constant dry weight (Wd). Finally, the percentage of water absorption (Wabsoprtion) and the remaining mass (Rmass) were determined according to Eq. (1) [31] and Eq. (2) [44], respectively.

$$W_{absorption} {(}wt\% {)} = \left( {\frac{{{(}W_w - W_d {)}}}{W_d }} \right) \times 100$$
(1)
$$R_{mass} {(}wt\% {)} = \left( {\frac{W_d }{{W_i }}} \right) \times 100$$
(2)

pH

The pH value of the PBS solution was measured using a pH meter (Thermosicentific Eutech Elite pH, Singapore) at the previously designated periods.

Mechanical Evaluation

The mechanical properties were tested utilizing a Universal Instron testing machine at room temperature following the Standard Test Method for Tensile Properties of Plastics, ASTM D638, see Fig. 1d. The samples were subjected to uniaxial tension at a constant displacement rate of 1 mm/min until the break. Three samples were analyzed for each blend or PLA; the result is the average of these measurements.

Scanning Electron Microscopy

Surface and cross-section morphology of the PLA and its blends were observed employing a scanning electron microscope, Hitachi SU3500, operated at 5.00 kV. Cross-section samples were prepared by cryo-fracture of the probes by immersion in liquid nitrogen. The samples were coated with a thin layer of gold prior to analysis.

Cytotoxicity: MTT Assay

In order to evalute the possible damage that EPS1, EPS2, and EPS3 can cause to cells, decreasing in vitro cell viability, an MTT assay was performed according to ISO-10993-5. A total of 10,000 MSC were seeded per well in a 96 well microplate. Cells were cultured at 37 °C in a CO2 atmosphere (5%) with supplemented DMEM F-12 medium. Following 72 h of culture, cells were exposed to PLA 5, PLA 15, EPS1, EPS2, and EPS3 by quadruple. Samples were cut in circular shapes with 6 mm diameter. Cells with no film exposure were taken as a negative control. The samples were previously sterilized for 15 min using UV light exposition on each side. After 24 h of exposition with the films, cell death was measured by MTT assay. Media were removed, and cells were incubated with 100 μL of the MTT solution (Cell Proliferation Kit I, Roche, Basilea, Switzerland) for 3 h. After incubation, 100 μL of isopropanol pH 3 was added per well to dissolve formazan crystals. According to kit instructions, the plate was read in a microplate absorbance reader (Agilent Bio Tek Cytation Instruments, WI, USA). The cell viability percentages were obtained with Eq. (3).

$$Sample viability \% = \left( {\frac{sample absorbance}{{negative control absorbance}}} \right) \times 100$$
(3)

An analysis of variance (ANOVA) and Tukey’s HSD (Honestly Significant Difference) test, with a confidence interval of 95%, were performed to determine significant differences between the control group and the tested samples.

Cytotoxicity: H&E Staining

The MSC were seeded per well in a 96-well microplate, a total of 10,000 MSC. Cells were cultured and exposed to films same as the MTT assay. A hematoxylin and eosin (H&E) staining was performed according to ISO-10993-5. The culture media was removed, and cells were washed with PBS solution (Thermo Fisher Scientific, Waltham, MA, USA) three times. The cells were incubated for 10 min at −20 °C with cold methanol initiating the fixation procedure. Subsequently, cells were rewashed with PBS three times to eliminate methanol residuals. For the stain, the incubated cells were placed in hematoxylin for 5 min. Then, these were washed (two cycles) with distilled water and HCl (diluted at 0.5% in ethanol). Immediately, eosin was placed for 5 min, and finally, the cells were distilled water washed. The morphology of the cells was analyzed in an optical microscope (Inverted microscope CKX41, Olympus, Shinjuku, Japan).

Hemolysis Test

A hemolysis test was performed following the Standard Practice for Assessment of Hemolytic Properties of Materials ASTM F756-17 to measure the impact of erythrocyte lysis that the EPS1, EPS2, EPS3, PLA5, and PLA15 films can produce. A trained health worker took 3 mL of a blood sample from a healthy donor into a heparin tube utilizing venipuncture (Becton Dickinson, Franklin Lakes, NJ, USA). It is important to note that biosecurity parameters were taken into account. Erythrocytes were separated by centrifugation (3000 rpm for 4 min) and washed thrice with PBS solution. A final solution of 1% v/v of erythrocytes in PBS was prepared.

Moreover, each film's circular pieces (6 mm diameter) were sterilized with UV light exposition for 15 min on each side. The films were incubated in microtubes, by quadruplicate, in 200 μL of the 1% v/v erythrocytes solution for 30 min, at 37 °C and 400 rpm. An erythrocytes solution (1% v/v) was taken as a negative control (NC), not exposed to films. An erythrocytes solution in distilled water (1% v/v) was taken as a positive control (PC). After incubating samples and controls, tubes were centrifuged at 14,000 rpm for 3 min. Hemoglobin released in the supernatant was measured by UV–Vis spectroscopy (NanoDrop, Thermo Fisher Scientific Waltham, MA, USA) at 415 nm. The percentage of hemolysis was obtained with Eq. (4).

$$Hemolysis \% = \frac{{{(}Sample_{abs} - NC_{abs} {)}}}{{{(}PC_{abs} - NC_{abs} {)}}} \times 100$$
(4)

Results and Discussion

The polymeric blends, namely EPS1, EPS2, and EPS3, have identical compositions; however, they differ in the mixing time and the order of the polymer’s addition. The mixing time for EPS1 is 5 min, comparable with neat PLA5. The EPS2 and EPS3 samples have a mixing time of 15 min compared to neat PLA15. As for the order of the polymer’s addition, EPS1 and EPS2 have identical methods; in contrast, EPS3 was formulated by mixing PLA and EVA for 10 min, and then SMMA was added, mixing them for 5 min completing 15 min.

Water Absorption and Remaining Mass

The impact of the hydrolytic degradation on EPS1, EPS2, and EPS3 blends was investigated. Since water absorption and remaining mass are intimately related to hydrolytic degradation were monitored for 168 days. Figure 2 illustrates hydrolytic degradation behavior regarding time (Fig. 2a). As noted for the first 7 days, the liquid absorption increased rapidly, Fig. 2a, reaching a maximum of 0.83% for the neat PLA15 sample. Subsequently, the absorption rate decreases until it increases after 28 days of immersion. At the end of 84 days, the EPS1 and EPS2 blends reported the lowest percentage of water absorption, with values of 0.78% and 0.79%, respectively. Nevertheless, the PLA15 sample reported the highest amount of water absorption, around 1.04% at 168 days (Fig. 2a).

Fig. 2
figure 2

Evolution of water absorption during hydrolytic degradation. a EPS3, b EPS2, c EPS1, d PLA15, and e PLA5

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Singla et al. [45] studied PLA/EVA blends and mentioned that the presence of EVA droplets in the PLA facilitates water diffusion in the PLA/EVA interface. Nevertheless, in this study, SMMA lies between PLA and EVA bubbles hindering the water diffusion inside the sample, according to our previous work [34].

Additionally, it is well known that water absorption in PLA accelerates hydrolysis due to the cleavage of ester bonds [46]. The degradation property and velocity have also been related to this phenomenon [47]. Liu et al. reported that PLLA scaffolds undergo a first period of degradation, where PLLA absorbs a volume of the PBS solution, allowing swelling and subsequent hydrolytic degradation of the lower molecular weight PLLA. In addition, the gradual increase in the amount of liquid absorbed generates defects and porosity [31].

The results of the remaining mass (Rmass) are illustrated in Fig. 3; according to Eq. 2. The general changes presented by the samples are insignificant, with a variation of ±1%. However, the comparison of polymer blends reveals that EPS1 and EPS2 blends recorded the lowest weight increase than EPS3. In our previous work [34], EPS1 and EPS2 morphologically exhibited a homogeneity in EVA microbubbles suggesting that longer processing time is needed, as well as adding SMMA at the commencement of mixing formulation. In contrast, the morphology of EVA for EPS3 is considerably heterogeneous due to the coalescence process, which leads to higher water absorption than EPS1 and EPS2. This result can be explained by the formation of EVA bubbles, inhibiting the water diffusion towards the inside of the test probe, thereby slowing the degradation rate. Wongwiwattana and Thomas [33] observed a similar behavior because most degradation reaction products remain inside the samples. Likewise, in this study, it is observed that the PLA15 sample gains weight in a more significant proportion, considered as an example at 84 days, a value around 0.22%, compared to the EPS blends. This behavior suggests that PLA undergoes a faster hydrolytic degradation rate than EPS mixtures because they contain EVA and SMMA.

Fig. 3
figure 3

Remaining mass of the samples. a EPS3, b EPS2, c EPS1, d PLA15, and e PLA5

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Additionally, there are studies of PLA degradability utilizing phosphate solutions at 37 °C with results different from this study; notwithstanding, the measurement methodology is distinct. Navarro et al. [32] reported no weight loss in PLA/PCL blends with a PLA content between 70 and 100%, even after 140 days of hydrolytic degradation, while blends with lower content of PLA lose a 3 wt% after 140 days. Nevertheless, they changed the PBS solution once a week. Wongwiwattana and Thomas [33] reported that blends with PLA do not degrade because their glass transition temperature is above 37 °C and is too low to cause a significant hydrolysis reaction since the reaction is limited to the diffusion of PBS into the samples. Likewise, it is reported that PLA commonly presents hydrolysis at temperatures higher than 50 °C [7, 48]. Navarro et al. [32] also mentioned that PLA degrades by chain scission across the hole polymer chains; however, they remained inside and did not disappear during the washing of the sample, originating no weight loss. Other investigations have reported that the degradation process of PLA is by hydrolytic degradation of the ester group, and the subproducts of this degradation can be assimilated [49, 50].

To summarize, the minor changes in the remaining mass of the EPS1, EPS2, and EPS3 blends can be explained due to the low rate of hydrolytic degradation occurring at 37 °C, and the PBS solution was not changed over time.

pH

The hydrolytic degradation of the PLA and EPS blends was investigated by measuring the pH of the PBS solution at different times, Fig. 4. The behavior of the physicochemical property pH is a decrease close to 0.4 until day 49. It is well known that the byproducts of the hydrolytic degradation of PLA are functional groups of carboxylic acid type; therefore, the pH diminishes [51, 52]. However, Gorrasi and Pantani [24] observed a pH increment on amorphous and semicrystalline PLA in the first 6 days, followed by a decrease depending on the D isomer content, suggesting an acid-based hydrolysis mechanism. In addition, hydrolysis degradation tests were carried out in distilled water with a pH of 5.7 and at 58 °C.

Fig. 4
figure 4

Evolution of pH during hydrolytic degradation. a EPS3, b EPS2, c EPS1, d PLA15, e PLA5

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The pH, at 84 days, presented a gradual increase, reaching a value of around 7.37 and 7.33 for the EPS2 blend and PLA15, respectively. In contrast, at 168 days, the pH of the PBS solutions was reduced. For example, the solution corresponding to the samples EPS1, EPS2, and EPS3 registered an acidity between 6.5 and 6.6. According to Gorrasi and Pantani [24], PLA hydrolysis can occur in three phases: (1) diffusion of H3O+ in the polymeric bulk, that is, an excess of OH– in the solution, which would justify the increase in pH; (2) hydrolysis reaction, and (3) counter diffusion of reaction products (carboxylic acids, alcohols, oligomers). Therefore, the observed changes in pH at day 84 could be attributed to the beginning of OH– production at this time of the hydrolytic degradation test.

Scanning Electron Microscopy

Physical degradation of polymers can develop in two ways: bulk or surface erosion. Bulk erosion is when degradation proceeds throughout the polymer matrix, while surface erosion is located only on the polymer surface [25, 53, 54]. Figure 5 illustrates the surface morphologies of PLA5, PLA15, EPS1, EPS2, and EPS3 samples at 0, 49, and 168 days of hydrolytic degradation. All micrographs of the samples at different times are observed in the supplementary material. Compared with the neat PLA5 at day 0, the samples of neat PLA5 at 49 and 168 days present a smoother surface (Fig. 5a–c). In addition, on day 168, slight flaking is observed, which can be attributed to the beginning of the polymer degradation. As can be observed, the neat PLA15 sample presents higher roughness (Fig. 5d–f), as well as the generation of small holes over its surface since day 49 (Fig. 5e), which is attributable to surface erosion. On day 168, these holes transform into cracks and are more prominent, suggesting higher degradation.

Fig. 5
figure 5

Surface scanning electron micrographs of PLA and EPS blends at (0), (49), and (168) days of exposure to PBS solution at 37 °C. a–c PLA5; d–f PLA15; g–i EPS1; j–l EPS2; m–o EPS3

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Moreover, the EPS blends show differences in surface morphology when compared regarding time, see Fig. 5g–o. For example, on day 49, EPS2 and EPS3 show holes on the surface; on the contrary, EPS1 has a smoother surface suggesting that this blend was degraded to a minimal grade. In the same way, it is observed at 168 days. These results can be explained by the processing conditions used to obtain the probes. The EPS1 blend has the shortest mixing time compared to the 15 min of EPS2 and EPS3. Hence, a mixing time of 5 min is necessary to decrease the hydrolytic degradation of the PLA blend.

In order to analyze the morphology of the cross-section during hydrolytic degradation, the probes were fractured with liquid nitrogen. Figure 6 illustrates the micrographs of PLA and EPS blends during hydrolytic degradation. The formation of holes is evident on day 168 for the neat PLA5 and PLA15 (Fig. 6a–f). Compared with neat PLA5, the PLA15 sample exhibits numerous pores evenly distributed throughout the probe, suggesting more bulk hydrolytic degradation. Zamani et al. [55] reported that degradation occurs at the surface when the degradation rate of the polymer is faster than the permeation rate of the solvent into the polymer matrix. Otherwise, degradation will take place inside the polymer matrix. Furthermore, Tsuji et al. [56] mention that the bulk erosion mechanism of PLA is dominant when hydrolyzed in PBS (pH 7.4).

Fig. 6
figure 6

Cross-section scanning electron micrographs of PLA and EPS blends at (0), (49), and (168) days of exposure to PBS solution at 37 °C. a–c PLA5; d–f PLA15; g–i EPS1; j–l EPS2; m–o EPS3

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Additionally, some authors mention that the hydrolytic degradation of amorphous poly (dl-lactic acid) proceeds heterogeneously and is faster inside than on the surface since a more significant contribution of autocatalysis occurs [57,58,59]. This behavior is consistent with our water absorption results, see Fig. 2. In previous work, a representative model of EPS blends was described, illustrating the role of each component in the ternary blend coexistence [34]. Concerning EPS1 (Fig. 6g–i), the small pores become evident after 49 days of degradation. In the case of EPS2 (Fig. 6j–l), fewer pores are observed on day 168, attributed to the homogeneity in size and distribution of EVA microbubbles generated during processing [6]. This homogeneity hinders the diffusion of the liquid inside the probe, decreasing the degradation rate of the bulk material.

Mechanical Evaluation

The mechanical properties of the PLA and EPS blends were evaluated during hydrolytic degradation, specifically elastic modulus (E), tensile strength (TS), and elongation at break. Figure 7 displays the E of the PLA5, PLA15, EPS1, EPS2, and EPS3 samples. In the EPS blends (Fig. 7a–c), at time 0, the values of E were similar, around 0.52 GPa. However, on day 49, the samples EPS1, EPS2, and EPS3 show a decrease in modulus E of 13%, 10%, and 6%, respectively. It is important to note that at 49 days was recorded that the mixtures began to retain saline liquid, which could be related to this decrease in E. In the case of PLA5, no significant affectation was observed in E throughout the test. However, the PLA15 sample consistently displayed lower modulus values than the neat PLA5 sample, indicating a more significant degradation acquired during the processing of PLA for a longer time, 15 min.

Fig. 7
figure 7

Young’s modulus behavior during hydrolytic degradation. a EPS3, b EPS2, c EPS1, d PLA15, e PLA5

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Figure 8 displays the TS in tension measured at different hydrolytic degradation times of all samples. TS represents the probe’s resistance to elongation [60]. All EPS blends report lower values of TS than PLA (Fig. 8a–e), attributed to the EVA content in the blend. TS presents more fluctuations for EPS blends during the hydrolytic degradation than PLA. These results could be associated with the heterogeneous morphology of EPS blends mainly caused by EVA bubbles.

Fig. 8
figure 8

Tensile strength during hydrolytic degradation. a EPS3, b EPS2, c EPS1, d PLA15, e PLA5

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Additionally, TS was more affected for PLA15 compared to day 168. In addition, the neat PLA15 sample presented a stress value of 27% lower than neat PLA5 at the end of the degradation test. This result could be due to a substantial amount of water absorption, which favors the degradation of the material and hence mechanically deteriorates PLA15 in a significant proportion favoring the sliding of the polymeric chains. However, compared with neat PLA5 on day 0, the neat PLA15 exhibits slightly lower TS due to a degradation of the polymer during the mixing and molding process. In particular, the neat PLA15 presented a lower viscosity than the neat PLA5 indicating a more significant degradation due to a shortening of polymer chains. The viscosity analysis was done in our previous work [34]. This shortening can also facilitate liquid diffusion into the material, thereby lowering mechanical performance. According to Rosli and coworkers [61], when the PLA is in an amorphous state, it allows a considerable diffusion of water inside the polymer, causing a faster degradation and, therefore, a decrease in mechanical properties.

The elongation at break of the PLA and EPS blends is displayed in Fig. 9. The EPS3 blend reported the lowest elongation percentage at time zero regarding the EPS blends, see Fig. 9a–c. As the degradation test elapsed, the elongation percentage had erratic behavior. However, on day 168, all blends reported lower elongation values than those reported on day 0. Likewise, the EPS3 sample had the lowest elongation on day 168, showing a more brittle behavior due to degradation. Furthermore, EPS2 displayed higher mechanical stability during hydrolytic degradation. This finding can be attributed to less bulk degradation, as observed by microscopy. Equally important is the processing conditions of this blend, especially 15 min of mixing all its components. On the other hand, the fluctuations in elongation at break could be attributed to differences in morphology acquired during sample processing. As previously observed by microscopy EPS blends present differences in the homogeneity in size and distribution of EVA microbubbles generated during processing, which are the main contributor to the elongation capacity of the blends. For some medical applications, such as drug delivery, surface degradation is more desirable because it permits the delivery of the drug to maintain the mechanical integrity of the material [62], as in the case of EPS2.

Fig. 9
figure 9

Elongation at break behavior during hydrolytic degradation. a EPS3, b EPS2, c EPS1, d PLA15, e PLA5

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The PLA15 sample reports appreciable elongation from 0 to 49 degradation days. This result can be attributed to the fact that, as previously mentioned, this sample suffered considerable degradation during processing. Therefore, its polymeric chains are smaller, which could favor movement when the sample is stressed, resulting in significant elongation. However, on days 84 and 168, PLA15 reported the lowest elongation caused by a higher rate of water absorption, i.e., a higher rate of degradation. Finally, PLA5 (Fig. 9e) did not present significant changes during hydrolytic degradation.

Vieira et al. [52] reported a negligible variation in Young’s modulus of the fibers, in PLA-PCL fibers, during the degradation process with saline phosphate solution at pH 7.4 and a temperature of 37 °C. Likewise, Zhou et al. [63] studied mixtures of polybutylene succinate with PLA subjected to hydrolytic degradation in a solution simulating human body fluids at a pH of 7.4 and a temperature of 37 °C. They reported no statistical difference in the tensile strength of PLA until about 6 months of degradation and reported 60% in mechanical strength, indicating that PLA experiences slower degradation compared to PBS. Additionally, they mention that elongation at break is the most sensitive index for monitoring the anchorages of the polymeric chains.

Possible applications are the manufacture of tendons and ligaments and as a replacement for the porous part of the bone (also called trabecular or spongy bone). Tensile strength from 4.4 to 660 MPa for ligaments and tendons of different parts of the human body was reported [64,65,66]. Regarding the maximum deformation and Young’s modulus of these connective tissues, values between 18 and 30% and 0.2 to 1.5 GPa were found, respectively [67, 68]. Specifically, in trabecular bones, which constitute the porous or spongy part of the bones, the modulus of elasticity ranges from 0.02 GPa to 0.5 GPa [69, 70].

Evaluation of the Cytotoxicity

Cytotoxicity tests evaluate cell damage, growth, and metabolism. MTT assay measures the metabolic activity of cells. The results of the MMT assay are shown in Fig. 10. The materials tested showed no statistical difference (p = 0.05) compared with the negative control, which can be considered that there is no evidence of damage in the cell metabolism. Therefore, EPS blends should not be considered dangerous substances.

Fig. 10
figure 10

MMT results of PLA and EPS blends

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The proliferation of mesenchymal cell growth in PLA and PCL nanofibrous scaffolds has previously been studied by Marei et al. [71]. They established that the cytotoxicity of both polymers was negligible because BMSC showed similar cell viability at day 3 compared to the control group.

To observe morphological changes in the cells due to their exposure to EPS blends and their comparison with neat PLA and negative control, H&E was performed. Figure 11 shows images from an optical microscope at 40× after 24 h of exposure. Cell morphology changes and cell confluence alterations were not observed in the evaluated samples. In addition, there are no evident changes in the cytoplasm or nucleus morphology. H&E results are comparable with MMT results, with both assays showing no evident cell damage.

Fig. 11
figure 11

Optical microscope images at 40× of H&E staining at 24 h. a Negative control, b PLA5, c PLA15, d EPS1, e EPS2, f EPS3

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Evaluation of the Hemolysis

Hemolysis assay involves exposing the material to erythrocytes recently extracted from the human body. Red blood cells are sensitive, and their membrane will lyse if the materials are incompatible. The percentage of hemolysis can be measured because the erythrocytes have hemoglobin inside them. All erythrocytes that have been lysed will release hemoglobin, which spectroscopy UV–Vis measure at 415 nm. Compatible materials will not lyse erythrocytes; therefore, no hemoglobin will be released into the medium.

Hemolysis assays are a reference to know if biomaterials present some toxicity levels. Biomaterials must not produce cellular damage to any components, including membranes. Membrane lysis can be measured by hemolysis assay in human erythrocytes. In this assay, hemoglobin is released at the medium when the outer membrane of the human erythrocytes is destroyed. Lysis produces the release of all the intracellular content from the red cells. The measurement of released hemoglobin estimates the number of erythrocytes destroyed. The Standard Practice analyzed the results for Assessment of Hemolytic Properties of Materials ASTM F756-08, which indicates that any material with a hemolysis rate of less than 5% is not considered high hemolytic [72]. All the tested samples are below 5% hemolysis, Fig. 12; therefore, none of the materials are considered hemolytic. However, some of them are above 2%, indicating they are low hemolytic, so this must be considered to propose a relevant medical application. However, the PLA5 also exceeded the value of 2% in two of its specimens. Since PLA is a material considered suitable for medical applications, if we take the value of this last material as a reference, all the samples could be used without problems in developing a medical device in terms of cytotoxicity.

Fig. 12
figure 12

Hemolysis results. a EPS3, b EPS2, c EPS1, d PLA15, e PLA5

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Conclusions

The hydrolytic degradation study demonstrated that the ternary samples reported less absorbed water than the neat PLA. This result suggests that the incorporation of EVA and SMMA into the PLA matrix retards its rate of hydrolytic degradation, providing the possibility of having control of this type of degradation. The highest hydrolytic degradation was observed in the PLA15 sample, suggesting that longer mixing time favors the PLA molecule's scissions and hence its hydrolytic degradation.

The elastic modulus and tensile strength did not register significant variations in any sample during the entire hydrolytic degradation test, except in the case of PLA15, which experimented a decrease in tensile strength from 51.09 MPa to 36.86 MPa from day 0 to 168, respectively. This finding could be due to a substantial amount of water absorption, which favors the degradation of the material and hence mechanically deteriorates PLA15 in a significant proportion favoring the sliding of the polymeric chains.

Concerning elongation at break, it presented fluctuations during the 168 days of testing owing to sample morphology and the processing conditions. In addition, this property was the most sensitive to hydrolytic degradation. Compared with EPS1 and EPS2, EPS3 showed the lowest elongation at break on day 168, with a value of 6.46%. In contrast, EPS2 had a stable mechanical performance during hydrolytic degradation. This behavior can be attributed to a more homogeneous morphology acquired, where all components were mixed for 15 min. Compared with the samples, PLA15 displayed the lowest elongation at break on days 84 and 168, 6.13%, and 6.28%, respectively, due to higher water absorption.

Biocompatibility tests showed that EVA/PLA/SMMA blends showed no statistical difference compared with the negative control, which makes them non-toxic. EPS blends presented good hemocompatibility with values of less than 5%, considered an acceptable limit.