Photomicrographs of in situ forming implant degraded in vitro (scale bar: 10 

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Context 1

... (3) assumes the absence of closed pores (all pores are interconnected), and a pore volume equal to the volume of water uptaken by the implant. This parameter refers to the water present not only inside the porous structure, but also inside the interstices between the polymer chains. For each implant we determined the weight average molecular weight ( M w ) of PLGA by size exclusion chromatography (SEC) to study the degradation, and performed scanning electron microscopy (SEM) to assess their morphology. SEM photomicrographs were analyzed with an image processing program (ImageJ 1.40g). The pore size of the spongy matrix was determined by measuring both the pore diameter (expressed in m) and the pore area (expressed in m 2 ). Adittionally, we determined the pore density (number of pores per unit of surface area, expressed in pores/mm 2 ) and the relative porous area (area occupied by pores/total area · 100). Volumes of the polymer solution corresponding to 75 mg of PLGA ( M 0 ) were injected subcutaneously into Wistar rats of 250-300 g for in situ implant formation. Rats were maintained in standard laboratory conditions. Implants were removed at different post injection times, dried under vacuum to constant weight ( M dry ), and then stored at -20 °C for further analysis. The study was conducted by duplicate. Each implant was assesed to determine: Mass Loss according to Eq. (1), M w by SEC, and morphology by SEM. In this case, besides the pore size, pore density , and relative porous area , we also measured the relative channel area (area occupied by channels/total area · 100). A portion of the affected skin was removed, fixed in formol buffer and included in paraffin for histopathological evaluation. Animals tolerated well the PLGA solutions, since the inflammatory response at the implant site was mild and reversible, similar to that reported by other authors (Tang and Singh, 2009). Data normality was evaluated on the basis of the Shapiro Wilks test. Results that presented a normal distribution were reported as mean value standard deviation (SD). Results that did not present a normal distribution were reported as median and interquartile range (IR= quartile 1 quartile 3). The nonparametric Kruskal-Wallis test and the Mann-Whitney test allowed to find differences between several independent samples or to compare between pairs of independent samples, respectively. Samples were considered significantly different when p <0.01. Fig. 1 shows the in vitro evolution of the Mass Loss , Water Content , Hollow Fraction , and M w of the implants. The Mass Loss rate was fairly constant during the assay, losing 85% of the mass at day 43. The Water uptake was important at the beginning of the experiment (77% at 1 day), and occurred mainly during the first three weeks where implants incorporated aproximately 90% of water. The time dependence of the Water Content rate ( r Water Content ) (inlet Fig. 1 b) clearly showed this result, where the r Water Content presented a gradual reduction until it reached a plateau. The Hollow Fraction against time is presented in Fig. 1 c. The increase observed involved the water gain and also the loss of mass in the polymer matrix, as it is defined by the Eq. (3). This may explain the similar behaviour observed between Water Content and Hollow Fraction (see Fig 1 b-c). It has been suggested not to use water to measure the pore volume of a copolymer when water swells the matrix (Yan et al., 2000), because part of the water tends to penetrate and relax the polymer matrix (Arifin et al., 2006). But in our definition of the Hollow Fraction we considered all the water gain, the one that occupies the porous structure and the one that swells the polymer. The PLGA implants presented a first period where the M w decreased (Fig. 1d) until it reached an asymptotic value around 3800 g / mol up to day 50 of the degradation. The Mass Loss and M w of the implants during the in vivo assay are shown in Fig. 2. Similarly to the results in vitro , the Mass Loss rate was almost constant during the assay, but it reached a 90% loss at day 18. At day 22 post injection, the mass of the implant was negligible. The half-life of in vivo implants was significantly shorter than that of in vitro implants, as reported in previous studies (Lu et al., 2000). The M w of PLGA decreased rapidly during the first days up to 4000 g / mol. This value, similar to that achieved in vitro , was maintained until day 12. On day 18, the M w decreased even more, reaching a value of 2000 g/mol. The fact that in vivo implants reached a much smaller M w value than in vitro would probably be due to the hydrolytic action of immune cells and/or enzymes, which cleaved the polymer chains in smaller oligomeric units (Tracy et al., 1999). Fig. 3 and Fig. 4 show the sponge-like structure of the in situ formed implants and the changes of the morphology during the degradation study. Both in vitro and in vivo implants presented a continuos porous matrix with a large number of interconnected pores. Besides the pores, in vivo implants presented channels in its entire matrix. The structure of these implants was more compact, probably due to the pressure exerted by the skin during the precipitation process. The structure of the pores changed during the degradation in vitro . Implants show spherical pores which size clearly increased from day 3 to 12 of the assay. At day 22 pores adquired a non-spherical shape, while after 33 days the amorphous structure of pores clearly showed an advanced erosional process (Fig. 3). On the in vivo assay, implants of day 3 and 12 post injection had similar structures. As degradation progressed, implants seemed to have a larger area occupied by channels and a greater ammount of very small pores. Table 1 shows the results of the photomicrographs analysis of in vitro and in vivo implants at different degradation times. The pore size of the in vitro implants -measured as pore area and pore diameter - significantly increased from day 3 to day 12 of degradation. At day 22, pores maintained their size but acquired non-spherical structures. Therefore, in that case we measured the area and not the diameter. Quantification at day 33 was not possible due to the high degree of erosion. The pore area of in vivo implants significantly decreased from day 3 to day 12 post injection, related to the formation of new smaller pores. This phenomenon woul probably be associated to a degradation process more than an erosional process. The pore area of the implants increased in vitro, while it decreased in vivo . Conversely, this difference is reflected in the pore density . This parameter decreased in vitro due to an increase in pore size, while it increased in vivo because of the appearance of large number of small pores. As degradation time progressed, relative porous area decreased both in vitro and in vivo . These differences are not statistically significant but they can be appreciated in Fig. 3., where pores start to fuse together and their structures are not easily differentiated. Furthermore, the area occupied by channels ( relative channel area ) in the in vivo implants increased from day 3 to day 12 of degradation. Channels of in vivo implants probably allowed physiological fluids to flow through the polymer matrix, maintaining the pH and preventing the accumulation of degradation products that could be exerting an autocatalytic effect. This physiological condition was simulated in vitro by making a rigurous control of the pH of the medium and replacing the buffer every time it dropped to 7.0. Otherwise, in vivo channels may facilitate degradation by increasing the contact surface area with aqueous fluids and allowing the entrance of hydrolytic enzymes and immune system cells to the entire matrix. Morphological changes observed in vitro during degradation were associated to an increase in pore size and a decrease in its density, while in vivo they were associated to the appearance of many new tinny pores and an increase of the area occupied by channels. Experimental results clearly show that an in vitro-in vivo correlation is not straightforward. Channels observed in the implants formed in vivo probably prevented the accumulation of degradation products that could lead to an autocatalytic effect. But on the other hand, they increased the contact with aqueous fluids, which in turns facilites the access of hydrolytic enzymes and immune system cells to the whole implant. This would explain the shorter degradation time of the implants degraded in vivo . These outcomes highlight the importance of the biological mechanisms involved in the degradation of PLGA implants. The author made this work in memory of Dr. Ricardo J.A. Grau and Dra. Ma. Inés Cabrera, who passed away while this paper was being done. Moreover, the author wish to thank to Agencia Nacional de Promoción Científica y Tecnológica (ANPCYT), Consejo Nacional de Investigaciones Científicas y Técnicas (CONICET), and Universidad Nacional del Litoral (UNL) of Argentina, for the financial support granted to this ...

Context 2

... (3) assumes the absence of closed pores (all pores are interconnected), and a pore volume equal to the volume of water uptaken by the implant. This parameter refers to the water present not only inside the porous structure, but also inside the interstices between the polymer chains. For each implant we determined the weight average molecular weight ( M w ) of PLGA by size exclusion chromatography (SEC) to study the degradation, and performed scanning electron microscopy (SEM) to assess their morphology. SEM photomicrographs were analyzed with an image processing program (ImageJ 1.40g). The pore size of the spongy matrix was determined by measuring both the pore diameter (expressed in m) and the pore area (expressed in m 2 ). Adittionally, we determined the pore density (number of pores per unit of surface area, expressed in pores/mm 2 ) and the relative porous area (area occupied by pores/total area · 100). Volumes of the polymer solution corresponding to 75 mg of PLGA ( M 0 ) were injected subcutaneously into Wistar rats of 250-300 g for in situ implant formation. Rats were maintained in standard laboratory conditions. Implants were removed at different post injection times, dried under vacuum to constant weight ( M dry ), and then stored at -20 °C for further analysis. The study was conducted by duplicate. Each implant was assesed to determine: Mass Loss according to Eq. (1), M w by SEC, and morphology by SEM. In this case, besides the pore size, pore density , and relative porous area , we also measured the relative channel area (area occupied by channels/total area · 100). A portion of the affected skin was removed, fixed in formol buffer and included in paraffin for histopathological evaluation. Animals tolerated well the PLGA solutions, since the inflammatory response at the implant site was mild and reversible, similar to that reported by other authors (Tang and Singh, 2009). Data normality was evaluated on the basis of the Shapiro Wilks test. Results that presented a normal distribution were reported as mean value standard deviation (SD). Results that did not present a normal distribution were reported as median and interquartile range (IR= quartile 1 quartile 3). The nonparametric Kruskal-Wallis test and the Mann-Whitney test allowed to find differences between several independent samples or to compare between pairs of independent samples, respectively. Samples were considered significantly different when p <0.01. Fig. 1 shows the in vitro evolution of the Mass Loss , Water Content , Hollow Fraction , and M w of the implants. The Mass Loss rate was fairly constant during the assay, losing 85% of the mass at day 43. The Water uptake was important at the beginning of the experiment (77% at 1 day), and occurred mainly during the first three weeks where implants incorporated aproximately 90% of water. The time dependence of the Water Content rate ( r Water Content ) (inlet Fig. 1 b) clearly showed this result, where the r Water Content presented a gradual reduction until it reached a plateau. The Hollow Fraction against time is presented in Fig. 1 c. The increase observed involved the water gain and also the loss of mass in the polymer matrix, as it is defined by the Eq. (3). This may explain the similar behaviour observed between Water Content and Hollow Fraction (see Fig 1 b-c). It has been suggested not to use water to measure the pore volume of a copolymer when water swells the matrix (Yan et al., 2000), because part of the water tends to penetrate and relax the polymer matrix (Arifin et al., 2006). But in our definition of the Hollow Fraction we considered all the water gain, the one that occupies the porous structure and the one that swells the polymer. The PLGA implants presented a first period where the M w decreased (Fig. 1d) until it reached an asymptotic value around 3800 g / mol up to day 50 of the degradation. The Mass Loss and M w of the implants during the in vivo assay are shown in Fig. 2. Similarly to the results in vitro , the Mass Loss rate was almost constant during the assay, but it reached a 90% loss at day 18. At day 22 post injection, the mass of the implant was negligible. The half-life of in vivo implants was significantly shorter than that of in vitro implants, as reported in previous studies (Lu et al., 2000). The M w of PLGA decreased rapidly during the first days up to 4000 g / mol. This value, similar to that achieved in vitro , was maintained until day 12. On day 18, the M w decreased even more, reaching a value of 2000 g/mol. The fact that in vivo implants reached a much smaller M w value than in vitro would probably be due to the hydrolytic action of immune cells and/or enzymes, which cleaved the polymer chains in smaller oligomeric units (Tracy et al., 1999). Fig. 3 and Fig. 4 show the sponge-like structure of the in situ formed implants and the changes of the morphology during the degradation study. Both in vitro and in vivo implants presented a continuos porous matrix with a large number of interconnected pores. Besides the pores, in vivo implants presented channels in its entire matrix. The structure of these implants was more compact, probably due to the pressure exerted by the skin during the precipitation process. The structure of the pores changed during the degradation in vitro . Implants show spherical pores which size clearly increased from day 3 to 12 of the assay. At day 22 pores adquired a non-spherical shape, while after 33 days the amorphous structure of pores clearly showed an advanced erosional process (Fig. 3). On the in vivo assay, implants of day 3 and 12 post injection had similar structures. As degradation progressed, implants seemed to have a larger area occupied by channels and a greater ammount of very small pores. Table 1 shows the results of the photomicrographs analysis of in vitro and in vivo implants at different degradation times. The pore size of the in vitro implants -measured as pore area and pore diameter - significantly increased from day 3 to day 12 of degradation. At day 22, pores maintained their size but acquired non-spherical structures. Therefore, in that case we measured the area and not the diameter. Quantification at day 33 was not possible due to the high degree of erosion. The pore area of in vivo implants significantly decreased from day 3 to day 12 post injection, related to the formation of new smaller pores. This phenomenon woul probably be associated to a degradation process more than an erosional process. The pore area of the implants increased in vitro, while it decreased in vivo . Conversely, this difference is reflected in the pore density . This parameter decreased in vitro due to an increase in pore size, while it increased in vivo because of the appearance of large number of small pores. As degradation time progressed, relative porous area decreased both in vitro and in vivo . These differences are not statistically significant but they can be appreciated in Fig. 3., where pores start to fuse together and their structures are not easily differentiated. Furthermore, the area occupied by channels ( relative channel area ) in the in vivo implants increased from day 3 to day 12 of degradation. Channels of in vivo implants probably allowed physiological fluids to flow through the polymer matrix, maintaining the pH and preventing the accumulation of degradation products that could be exerting an autocatalytic effect. This physiological condition was simulated in vitro by making a rigurous control of the pH of the medium and replacing the buffer every time it dropped to 7.0. Otherwise, in vivo channels may facilitate degradation by increasing the contact surface area with aqueous fluids and allowing the entrance of hydrolytic enzymes and immune system cells to the entire matrix. Morphological changes observed in vitro during degradation were associated to an increase in pore size and a decrease in its density, while in vivo they were associated to the appearance of many new tinny pores and an increase of the area occupied by channels. Experimental results clearly show that an in vitro-in vivo correlation is not straightforward. Channels observed in the implants formed in vivo probably prevented the accumulation of degradation products that could lead to an autocatalytic effect. But on the other hand, they increased the contact with aqueous fluids, which in turns facilites the access of hydrolytic enzymes and immune system cells to the whole implant. This would explain the shorter degradation time of the implants degraded in vivo . These outcomes highlight the importance of the biological mechanisms involved in the degradation of PLGA implants. The author made this work in memory of Dr. Ricardo J.A. Grau and Dra. Ma. Inés Cabrera, who passed away while this paper was being done. Moreover, the author wish to thank to Agencia Nacional de Promoción Científica y Tecnológica (ANPCYT), Consejo Nacional de Investigaciones Científicas y Técnicas (CONICET), and Universidad Nacional del Litoral (UNL) of Argentina, for the financial support granted to this ...

Context 3

... (3) assumes the absence of closed pores (all pores are interconnected), and a pore volume equal to the volume of water uptaken by the implant. This parameter refers to the water present not only inside the porous structure, but also inside the interstices between the polymer chains. For each implant we determined the weight average molecular weight ( M w ) of PLGA by size exclusion chromatography (SEC) to study the degradation, and performed scanning electron microscopy (SEM) to assess their morphology. SEM photomicrographs were analyzed with an image processing program (ImageJ 1.40g). The pore size of the spongy matrix was determined by measuring both the pore diameter (expressed in m) and the pore area (expressed in m 2 ). Adittionally, we determined the pore density (number of pores per unit of surface area, expressed in pores/mm 2 ) and the relative porous area (area occupied by pores/total area · 100). Volumes of the polymer solution corresponding to 75 mg of PLGA ( M 0 ) were injected subcutaneously into Wistar rats of 250-300 g for in situ implant formation. Rats were maintained in standard laboratory conditions. Implants were removed at different post injection times, dried under vacuum to constant weight ( M dry ), and then stored at -20 °C for further analysis. The study was conducted by duplicate. Each implant was assesed to determine: Mass Loss according to Eq. (1), M w by SEC, and morphology by SEM. In this case, besides the pore size, pore density , and relative porous area , we also measured the relative channel area (area occupied by channels/total area · 100). A portion of the affected skin was removed, fixed in formol buffer and included in paraffin for histopathological evaluation. Animals tolerated well the PLGA solutions, since the inflammatory response at the implant site was mild and reversible, similar to that reported by other authors (Tang and Singh, 2009). Data normality was evaluated on the basis of the Shapiro Wilks test. Results that presented a normal distribution were reported as mean value standard deviation (SD). Results that did not present a normal distribution were reported as median and interquartile range (IR= quartile 1 quartile 3). The nonparametric Kruskal-Wallis test and the Mann-Whitney test allowed to find differences between several independent samples or to compare between pairs of independent samples, respectively. Samples were considered significantly different when p <0.01. Fig. 1 shows the in vitro evolution of the Mass Loss , Water Content , Hollow Fraction , and M w of the implants. The Mass Loss rate was fairly constant during the assay, losing 85% of the mass at day 43. The Water uptake was important at the beginning of the experiment (77% at 1 day), and occurred mainly during the first three weeks where implants incorporated aproximately 90% of water. The time dependence of the Water Content rate ( r Water Content ) (inlet Fig. 1 b) clearly showed this result, where the r Water Content presented a gradual reduction until it reached a plateau. The Hollow Fraction against time is presented in Fig. 1 c. The increase observed involved the water gain and also the loss of mass in the polymer matrix, as it is defined by the Eq. (3). This may explain the similar behaviour observed between Water Content and Hollow Fraction (see Fig 1 b-c). It has been suggested not to use water to measure the pore volume of a copolymer when water swells the matrix (Yan et al., 2000), because part of the water tends to penetrate and relax the polymer matrix (Arifin et al., 2006). But in our definition of the Hollow Fraction we considered all the water gain, the one that occupies the porous structure and the one that swells the polymer. The PLGA implants presented a first period where the M w decreased (Fig. 1d) until it reached an asymptotic value around 3800 g / mol up to day 50 of the degradation. The Mass Loss and M w of the implants during the in vivo assay are shown in Fig. 2. Similarly to the results in vitro , the Mass Loss rate was almost constant during the assay, but it reached a 90% loss at day 18. At day 22 post injection, the mass of the implant was negligible. The half-life of in vivo implants was significantly shorter than that of in vitro implants, as reported in previous studies (Lu et al., 2000). The M w of PLGA decreased rapidly during the first days up to 4000 g / mol. This value, similar to that achieved in vitro , was maintained until day 12. On day 18, the M w decreased even more, reaching a value of 2000 g/mol. The fact that in vivo implants reached a much smaller M w value than in vitro would probably be due to the hydrolytic action of immune cells and/or enzymes, which cleaved the polymer chains in smaller oligomeric units (Tracy et al., 1999). Fig. 3 and Fig. 4 show the sponge-like structure of the in situ formed implants and the changes of the morphology during the degradation study. Both in vitro and in vivo implants presented a continuos porous matrix with a large number of interconnected pores. Besides the pores, in vivo implants presented channels in its entire matrix. The structure of these implants was more compact, probably due to the pressure exerted by the skin during the precipitation process. The structure of the pores changed during the degradation in vitro . Implants show spherical pores which size clearly increased from day 3 to 12 of the assay. At day 22 pores adquired a non-spherical shape, while after 33 days the amorphous structure of pores clearly showed an advanced erosional process (Fig. 3). On the in vivo assay, implants of day 3 and 12 post injection had similar structures. As degradation progressed, implants seemed to have a larger area occupied by channels and a greater ammount of very small pores. Table 1 shows the results of the photomicrographs analysis of in vitro and in vivo implants at different degradation times. The pore size of the in vitro implants -measured as pore area and pore diameter - significantly increased from day 3 to day 12 of degradation. At day 22, pores maintained their size but acquired non-spherical structures. Therefore, in that case we measured the area and not the diameter. Quantification at day 33 was not possible due to the high degree of erosion. The pore area of in vivo implants significantly decreased from day 3 to day 12 post injection, related to the formation of new smaller pores. This phenomenon woul probably be associated to a degradation process more than an erosional process. The pore area of the implants increased in vitro, while it decreased in vivo . Conversely, this difference is reflected in the pore density . This parameter decreased in vitro due to an increase in pore size, while it increased in vivo because of the appearance of large number of small pores. As degradation time progressed, relative porous area decreased both in vitro and in vivo . These differences are not statistically significant but they can be appreciated in Fig. 3., where pores start to fuse together and their structures are not easily differentiated. Furthermore, the area occupied by channels ( relative channel area ) in the in vivo implants increased from day 3 to day 12 of degradation. Channels of in vivo implants probably allowed physiological fluids to flow through the polymer matrix, maintaining the pH and preventing the accumulation of degradation products that could be exerting an autocatalytic effect. This physiological condition was simulated in vitro by making a rigurous control of the pH of the medium and replacing the buffer every time it dropped to 7.0. Otherwise, in vivo channels may facilitate degradation by increasing the contact surface area with aqueous fluids and allowing the entrance of hydrolytic enzymes and immune system cells to the entire matrix. Morphological changes observed in vitro during degradation were associated to an increase in pore size and a decrease in its density, while in vivo they were associated to the appearance of many new tinny pores and an increase of the area occupied by channels. Experimental results clearly show that an in vitro-in vivo correlation is not straightforward. Channels observed in the implants formed in vivo probably prevented the accumulation of degradation products that could lead to an autocatalytic effect. But on the other hand, they increased the contact with aqueous fluids, which in turns facilites the access of hydrolytic enzymes and immune system cells to the whole implant. This would explain the shorter degradation time of the implants degraded in vivo . These outcomes highlight the importance of the biological mechanisms involved in the degradation of PLGA implants. The author made this work in memory of Dr. Ricardo J.A. Grau and Dra. Ma. Inés Cabrera, who passed away while this paper was being done. Moreover, the author wish to thank to Agencia Nacional de Promoción Científica y Tecnológica (ANPCYT), Consejo Nacional de Investigaciones Científicas y Técnicas (CONICET), and Universidad Nacional del Litoral (UNL) of Argentina, for the financial support granted to this ...