Role of pore size and morphology in musculo-skeletal tissue regeneration

Materials Science and Engineering C 61 (2016) 922–939

Contents lists available at ScienceDirect

Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec

Review

Role of pore size and morphology in musculo-skeletal tissue regeneration Roman A. Perez a,b,⁎,1, Gemma Mestres c,1 a b c

Department of Nanobiomedical Science & BK21 PLUS NBM Global Research Center for Regenerative Medicine, Dankook University, Cheonan 330-714, Republic of Korea Institute of Tissue Regeneration Engineering (ITREN), Dankook University, Cheonan 330-714, Republic of Korea Department of Engineering Sciences, Uppsala University, Box 534, 751 21 Uppsala, Sweden

a r t i c l e

i n f o

Article history: Received 1 May 2015 Received in revised form 23 December 2015 Accepted 28 December 2015 Available online 31 December 2015 Keywords: Microporosity Macroporosity Porosity Pore size Scaffold Bone regeneration Microstructure Rapid prototyping

a b s t r a c t Biomaterials in the form of scaffolds hold great promise in the regeneration of diseased tissues. The scaffolds stimulate cellular adhesion, proliferation and differentiation. While the scaffold composition will dictate their biocompatibility, their porosity plays a key role in allowing proper cell penetration, nutrient diffusion as well as bone ingrowth. Porous scaffolds are processed with the help of a wide variety of techniques. Designing scaffolds with the appropriate porosity is a complex issue since this may jeopardize other physico-chemical properties. From a macroscopic point of view, parameters such as the overall architecture, pore morphology, interconnectivity and pore size distribution, have unique roles in allowing bone ingrowth to take place. From a microscopic perspective, the adsorption and retention of proteins in the microporosities of the material will dictate the subsequent cell adhesion. Therefore, the microstructure of the substrate can determine cell proliferation as well as the expression of specific osteogenic genes. This review aims at discussing the effect of micro- and macroporosity on the physicochemical and biological properties of scaffolds for musculo-skeletal tissue regeneration. © 2016 Elsevier B.V. All rights reserved.

Contents 1. 2. 3.

Introduction . . . . . . . . . . . . . . . . . . Types of porosity . . . . . . . . . . . . . . . . Macroporosity . . . . . . . . . . . . . . . . . 3.1. Tuning of mechanical properties . . . . . 3.1.1. Randomly distributed pores . . . 3.1.2. Computer designed pores . . . . 3.1.3. Aligned pores . . . . . . . . . . 3.1.4. Hierarchical pores . . . . . . . . 3.2. Biological properties (in vitro) . . . . . . . 3.2.1. Pore size . . . . . . . . . . . . 3.2.2. Interconnections and permeability 3.2.3. Pore morphology . . . . . . . . 3.2.4. Pore size distribution . . . . . . 3.3. Preclinical assays (in vivo) . . . . . . . . 3.3.1. Pore size and pore size distribution 3.3.2. Interconnections and permeability 3.3.3. Pore morphology . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

. . . . . . . . . . . . . . . . .

923 923 923 924 924 924 925 926 926 926 928 928 929 929 929 930 930

Abbreviations: ALP, alkaline phosphatase; BCP, biphasic calcium phosphate; BMP, bone morphogenetic protein; BMSC, bone marrow stromal cell; BSA, bovine serum albumin; CDHA, calcium deficient hydroxyapatite; CPC, calcium phosphate cement; μCT, X-ray micro-computed tomography; GAG, glycosaminoglycan; HA, hydroxyapatite; hBDC, human bone-derived cell; hBMSC, human bone marrow stromal cell; MSC, mesenchymal stem cell; PCL, poly Ɛ-caprolactone; PDLLA, poly-DL-lactic acid; PLLA, poly-(L-lactic acid); PLGA, poly-(lactic-co-glycolic acid); PVA, polyvinyl alcohol; SBF, simulated body fluid; SEM, scanning electron microscopy. ⁎ Corresponding author at: Department of Nanobiomedical Science & BK21 PLUS NBM Global Research Center for Regenerative Medicine, Dankook University, Cheonan 330-714, Republic of Korea. E-mail address: [email protected] (R.A. Perez). 1 Both authors have contributed equally in the manuscript.

http://dx.doi.org/10.1016/j.msec.2015.12.087 0928-4931/© 2016 Elsevier B.V. All rights reserved.

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

3.4. Macroporosity overview . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Microporosity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1. Tailoring osteoinductivity through microporosity — the importance of protein interaction with microporous substrates 4.2. Micropores in drug delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.3. Cell behavior on microporous materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.4. Preclinical assays (in vivo) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.4.1. Osteoinduction (subcutaneous models) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.4.2. Osteo conduction (bone and osteochondral defects) . . . . . . . . . . . . . . . . . . . . . . . . . . 5. Concluding remarks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.

1. Introduction Recent advances in tissue engineering have emphasized the need to properly design scaffolds to allow cells to attach, migrate, proliferate and differentiate [1,2]. While the composition and surface chemistry of scaffolds dictate the ability of cells to initially attach, the morphology of the scaffold plays a key role in controlling their ability to migrate [3,4]. Besides allowing cell penetration, a proper architecture of the scaffold will allow nutrients and oxygen to flow into it as well as to remove waste produced by the cells to increase cell survival and hence to regenerate tissue [2,3]. Thus, scaffolds need to be designed to present enough porosity, which not only has to have pores big enough to allocate cells, but also needs to present interconnections to allow cell migration between the different pores. Nevertheless, the pores not only play a significant role in allowing cell penetration and migration, but also significantly influence the physical properties of the scaffold [5–7]. For instance, the increase in the porosity is known to exponentially decrease the mechanical properties, whereas on the other hand, the permeability can be largely increased with increased porosity [2,8]. In order to optimize scaffolds for tissue engineering, all these parameters need to be balanced to guide proper tissue regeneration. The total porosity, the pore size and pore size distribution as well as the pore morphology are some key parameters that play a critical role in balancing the physical and biological properties of the scaffolds. Furthermore, these properties need to be balanced with the degradation of the scaffold. Ideally the scaffold should degrade at the same time as new natural tissue is being formed, which growth may be stimulated by the biomaterial [2,9]. The porosity is known to increase the ability of fluids to penetrate the structure and therefore enhances the degradation [4]. While pore sizes and pore interconnections in the range of hundreds of microns are relevant for cells to migrate and proliferate, pore sizes in a smaller range also play pivotal roles in tissue engineering [10,11]. These pores are usually few microns in size and are involved mainly on the initial adsorption of proteins on the surface of the materials. Cells interact with biomaterials through cell–protein interactions through the transmembrane proteins. Therefore, it is believed that the increase in protein concentration may significantly affect cell fate [12–15]. Besides the ability to adsorb proteins, these small sized pores are also known to allow the regulation of cell behavior, playing key roles in directing stem cell fate. The scope of this review is to provide a detailed description on how the different pore sizes and morphologies may affect the overall in vitro and in vivo tissue regeneration. At the same time, the review will discuss how the change in porosity or pore size affects other physical parameters that may also play important aspects in the overall bone regeneration such as the mechanical properties. While there are many types of scaffolds made of different materials and compositions, we do not aim to describe the different sources and the differences among them, but rather describe studies that have been able to analyze specific parameters of the scaffolds while maintaining other parameters constant. For this purpose, polymeric based materials are of great ease to

923

. . . . . . . . . . .

. . . . . . . . . . .

. . . . . . . . . . .

. . . . . . . . . . .

. . . . . . . . . . .

. . . . . . . . . . .

. . . . . . . . . . .

. . . . . . . . . . .

. . . . . . . . . . .

. . . . . . . . . . .

931 932 932 933 934 935 935 935 936 936 936

work since their mouldability allow to fully control their morphology and porosity as desired. While the porosity of ceramics can be controlled to a lower extent due to the high sintering temperatures, metallic materials will be overlooked as scaffolds in this review due to their inherent low biological properties as well as the possible release of ions that may cause severe adverse reactions [16,17]. For this purpose, the review will describe works mainly on polymeric materials as well as some ceramic materials that have allowed performing comparative studies. 2. Types of porosity Porosity is the quantification of void spaces within a material. The most common methods to measure the porosity are mercury intrusion porosimeter, capillarity and permeability methods. The main advantage of the mercury porosimeter is that it allows quantifying the pore size distribution of the pores' neck (detection limit of 0.06 μm) by incorporating mercury into the scaffolds through the use of pressure. The pore size distribution can also be obtained below this size by means of nitrogen adsorption. Other measurement techniques are based on imaging methods, such as micro computed tomography (μCT), scanning electron microscopy (SEM) as well as atomic force microscopy (AFM). These imaging techniques can reach high resolutions. For example, AFM can resolve pores as small as 500 nm, although the area of sample that can be studied is too small to obtain pore size distributions. Besides the porosity of the samples, the pore size distribution is of great importance as well. The porosities of dense materials are classified in three different types according to the IUPAC: micropores (b 2 nm), mesopores (2–50 nm) and macropores (N50 nm). Nevertheless, for tissue engineering, it is commonly used a slightly different description of the pore sizes. In this sense, we will adopt in this review the nomenclature used by biomaterial scientists to describe the pore sizes of scaffolds, which classifies pore sizes as macropores (N50 μm) and micropores (b50 μm). Therefore, we will only distinguish two different types of pores throughout the review and will not consider the pore ranges established by IUPAC. 3. Macroporosity Scaffold macroporosity plays a critical role in the regeneration of damaged tissues. Macroporosity is aimed to allow cell penetration, which may then trigger the integration with the host tissue and increase the chances for key processes to take place (e.g. blood vessel ingrowth). The optimum pore size for scaffolds lies in the range between 100 and 400 μm [2,18,19]. Nevertheless, it is still unclear whether scaffolds with, for instance, homogenous pore size distribution perform more efficiently than scaffolds with heterogeneous pores. These macroporosities can be obtained by a wide variety of techniques that include freeze drying, solvent casting, rapid prototyping or laser sintering, among others [2,20,21]. The fabrication methods will determine the pore morphologies and pore size, and hence need to be carefully chosen. Scaffold fabrication techniques are indicated and shortly described in Table 1. For further details about each of these

924

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

Table 1 Main fabrication techniques of scaffolds for bioactive glass, ceramic and polymeric materials. Class of biomaterials

Method/technique

Description of the method

References

Bioactive glass, ceramic

Chemical foaming method

[24]

Bioactive glass

Sol–gel processing

Bioactive glass, ceramic, polymer, ceramic/polymer Bioactive glass, ceramic

Sintering/Thermal bonding

A chemical foaming agent is mixed with a ceramic. Pyrolysis produces gas that foams the ceramic A sol is formed and condensed to form a gel. The gel is aged, dried and sintered to form scaffolds Particles are packed and thermally bonded a in a mold with the desired geometry A foam is immersed in the desired ceramic suspension to coat the foam struts. After drying, the foam and organic binders are burned out and the glass struts are sintered A porogen is added into a ceramic slurry. After scaffold fabrication, these are immersed in aqueous solution to leach the porogen. Colloidally-stable suspension of particles are frozen quickly and afterwards sublimated under cold temperatures under vacuum. The porous constructs are sintered. Air is incorporated into a ceramic suspension which is then set in order to create a structure of air bubbles Directional freezing of the suspensions allows the formation of porous scaffolds with an oriented microstructure. High voltage is applied to produce fibers from a liquid, and the fibers are used to build a scaffold Objects are manufactured layer-by-layer using the information from a software file

[32]

Foam/sponge replication/Sacrificial template method

Bioactive glass, ceramic, polymer, ceramic/polymer Bioactive glass, ceramic, polymer, ceramic/polymer

Porogen leaching (salt, paraffin, ice, gelatin, sugar, etc.) Freeze drying technique

Ceramic, polymer

Direct foaming method

Polymer

Temperature gradient-guided thermal-induced phase-separation technique Electrospinning

Polymer Bioactive glass, ceramic, polymer/glass, polymer

Rapid prototype technology

techniques the authors invite the readers to read key review papers on this topic [2,21–23]. 3.1. Tuning of mechanical properties The porosity of scaffolds is known to be detrimental for the mechanical properties, which tend to decrease exponentially as the porosity is increased [2,8]. Balancing the levels of porosity with the mechanical properties to properly guide tissue regeneration has become an issue of extensive research. While a classical method to control the porosity in ceramic materials has been the sintering temperature [8,34–36], the appearance of new and more sophisticated techniques for processing of scaffolds (e.g. 3D printing) has allowed an extensive control on the design that has significantly improved the mechanical properties. In this section we aim to address the possible systems in which the mechanical properties can be tuned to match that of musculo-skeletal tissues through an accurate design of pore size, porosity or pore interconnections. It is noteworthy that classical methods, which allow randomly distributed pores, have also evolved to allow better tuning of the mechanical properties. 3.1.1. Randomly distributed pores A vast number of scaffolds generated for tissue engineering purposes have been prepared with randomly distributed pores. These scaffolds were mainly designed to conceive porosity within the structure, without taking into account any pore characteristics. This is a common property for polymeric and ceramic scaffolds. For polymeric materials, salt leaching is an easy system to control the porosity of the scaffolds by simply changing the concentration and the size of salt particles. Synthetic polymeric scaffolds, mainly PLA or PGA and their derivates, are commonly prepared through salt leaching technique and in general show that smaller pores (100–200 μm) give rise to significantly lower elastic modulus compared to bigger pores (250–350 and 420–500 μm) [37]. Modifying the polymer used for the scaffold fabrication is another possibility to control the mechanical properties for scaffolds with identical porosity, showing for instance that PLA presents higher mechanical properties than PCL [38]. The use of porogen with different morphologies, mainly cubic and spherical, also shows effects on the structure and mechanical properties of these scaffolds, showing better mechanical properties for spherical pores [39]. Other polymeric scaffolds are those produced by supercritical CO2 foaming, which can easily vary the parameters to obtain different pores and interconnections to tune

[25] [26] [27]

[28] [29]

[30] [31]

[33]

the mechanical properties. For instance, by varying the molecular weight of PLA and the depressuration rates, scaffold morphology can easily be controlled, showing that high molecular weights with high depressuration rates provide optimum elastic modulus. This is mainly achieved through the production of scaffolds with homogenous pore distributions with higher material density [40]. While polymeric based scaffolds use techniques such as salt leaching or freeze-drying, ceramic materials commonly use sponge replication techniques. Glass–ceramic scaffolds with cancellous bone like structure, specifically with 56 vol.% total porosity and pore sizes ranged between 100 and 500 μm, have been produced by sponge replication. Taking into account their morphology, the mechanical properties of these structures were high showing a compressive strength of 18 MPa and elastic modulus around 380 MPa [41]. The same research group developed a compressive-tensile model to be able to describe porosity–strength relationships. The model allows predicting the mechanical properties of a scaffold with known porosity or, in the other way around, predicting the porosity using the information from the mechanical properties. This tool can be used to do custom-made scaffolds with the desired balance of porosity and mechanical properties [42]. Another model designed by Dejaco et al. evaluated how local stress increases chances of cracking, which in turn cause a decrease of the global stiffness of the structure [43]. Microcracking suffered by bone inside the body has been suggested to also stimulate the activity of bone cells [44]. 3.1.2. Computer designed pores The design of more complex and sophisticated scaffolds with unique pore morphology has allowed tailoring the pore morphology to decrease the mechanical properties to a lower extent. In this sense, rapid prototyping and 3D printing are powerful tools to perform systematic studies on the effect of porosities through an extensive control over pore architecture [45–49]. Scaffolds have been designed with different pore morphologies and sizes, presenting a homogenous/heterogeneous, monomodal/bimodal and random/organized pore distribution. All these parameters have a profound effect not only in the mechanical properties, but also in the biological properties of the material. All pore characteristics are relevant and their conjunction determines the mechanical properties of the scaffold. The effect of the pore morphology was evaluated by Lee et al. Synthetic polymeric matrixes based scaffolds were prepared by rapid prototyping, presenting strut morphology in lattice (Fig. 1a), staggered (Fig. 1b) or diagonal (Fig. 1c). Fig. 1d shows that the compressive strength increases as the architecture

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

925

Fig. 1. SEM images of scaffolds with different strut morphologies: a) lattice-, b) stagger-, and c) diagonal-type scaffold, and d) stress vs. strain plot for each scaffold (image from Lee et al. [45]).

becomes more complex, which is lattice b stagger b diagonal, having values of 6.05, 7.43 and 9.81 MPa, respectively [45]. Another work described similar scaffold morphologies, comparing PLGA lattice-like scaffolds versus stagger-like scaffolds with a displaced double layer design [46]. Due to the double layer displaced design, the compressive modulus decreased from approximately 90 MPa for the lattice-like to 28 MPa for the stagger-like scaffolds. The authors attributed the decrease of the mechanical properties to the separation between the layers, which lead to pore sizes that are too big to properly withstand mechanical stress [46]. Nevertheless, it is considered that the displaced morphology can be beneficial for the mechanical properties. In order to proof this concept, another work studied the effect of different offsets created in rapid prototyped PCL based scaffolds (Fig. 2) [47]. The results showed that as the offsets became higher, there was an increase in the elastic modulus, which was associated to the formation of smaller pores, despite the overall porosity in the different scaffolds was the same. This suggests that the presence of uniform small pores allows for enhanced mechanical properties. In a similar direction but a slightly different approach, a study compared homogenous versus gradient pore size distribution scaffolds (Fig. 3a) to determine which type of pore size distribution has better mechanical properties. Two different pore patterns were prepared for a homogenous (Homog 1 and Homog 2) and gradient-like scaffolds (Grad 1 and Grad 2). The pore size of Homog 1 and 2 was 750 μm and 100 μm, respectively. On the other hand, Grad 1 presented bigger sizes pores on the surface and the bottom of the scaffold (e.g. 750 μm), which decreased in size up to 100 μm towards the center of the scaffold. Opposite morphology was found for the Grad 2, which presented the bigger pore sizes in the center of the scaffolds and smaller pore sizes in the surface and the bottom of the scaffolds. The results showed that the strain–stress curves (Fig. 3b) presented very similar patterns for the gradient scaffolds. It was also revealed that the Homog 2 scaffolds with smaller pore sizes presented the highest supported stress, whereas the bigger homogenous pores presented the lowest values. This was

then correlated with higher Young modulus for the Homog 1 scaffolds (Fig. 3c) but with lower recovery rates (Fig. 3d) [50]. Therefore, it could be concluded that rather homogenous pores could enhance the mechanical properties, although gradient scaffolds can tailor their pore size and distribution in a more complex manner that can eventually lead to enhanced mechanical properties compared to homogenous pores. 3.1.3. Aligned pores Since bone can be considered as a hierarchical anisotropic composite material with high compressive loading on the z axis, a mimicking strategy consists on designing scaffolds with architectures similar to those of Haversian channels which may enhance the mechanical properties in the z axis. A common strategy to prepare these types of scaffolds is unidirectional freezing, which consists on controlling the ice crystal orientation during the freezing of a water based polymer or ceramic slurry. For instance, Wu et al. studied the mechanical properties of poly(ethylene glycol) (PEG) gels with anisotropic pores. The pores were introduced by unidirectional freezing and cryopolymerization technique. The mechanical properties were shown to be considerably anisotropic, having a compressive stress of 15 kPa and 18% strain when the force was applied parallel to the scaffold, and a compressive stress of around 3.5 kPa and 50% strain when the force was applied perpendicular to the scaffold [51]. Interestingly, the pores can be designed to be aligned as well in the y axis. In this sense, a recent work compared chitosan-gelatin scaffolds with randomly or aligned (vertical and horizontal) pores. The results showed that the mechanical properties were lowest for the scaffolds with horizontal aligned pores and maximum for the scaffolds with aligned pores in the vertical direction [52]. The pore alignment has not only shown to have direct effect on the mechanical properties of the scaffold, but also showed different interaction with biological tissues, and hence changed the mechanical properties of tissue-scaffold construct at the site of implantation. For instance, oriented or non-oriented scaffolds of bovine articular cartilage ECM were

926

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

scaffolds, which would provide support for compression in the vertical direction [53]. 3.1.4. Hierarchical pores Of special interest are the scaffolds that are able to present a bimodal pore size distribution, mainly having macropores for cells penetration and micropores for protein adsorption/delivery. Interestingly, the micropores also present some role in the overall mechanical properties of scaffolds. For instance, HA scaffolds were prepared with two different micropore sizes, mainly 5.96 and 16.2 μm, showing higher bending and compression strength in the presence of smaller micropores [54]. Microporous (54–80 vol.%) bioactive glass scaffolds with hierarchical porosity have been recently developed with the aim to satisfy the macroporosity needed for bone ingrowth while maintaining adequate mechanical properties (12–20 MPa). [48,55,56]. 3.2. Biological properties (in vitro) From a biological point of view, the macroporosity plays a key role in allowing enough cells to penetrate and remain within the scaffold boundaries. Several architectonic characteristics of the scaffold, such as pore size, interconnections, morphology and pore distribution, have shown to not only have an effect on their ability of cells to colonize the scaffolds but also on their ability to proliferate and differentiate.

Fig. 2. Surface and cross-sectional SEM images of PCL–β-TCP offset scaffolds fabricated using the melt-plotting system. Scale bars represent 200 μm (2nd and 4th row) and 500 μm (1st and 3rd row) (image from Yeo et al. [47]).

implanted in vivo in the dorsa of nude mice. After 2 and 4 weeks the tissue-scaffold construct were harvested to evaluate the mechanical properties showing that the value of the Young's modulus was around the double for the oriented scaffolds than for the non-oriented scaffolds. This was attributed to the alignment of collagen bundled in the oriented

3.2.1. Pore size Pore size has been profoundly studied in order to identify which is the optimum range for bone regeneration. A recent work showed the effect of pore size (100, 200, 350 and 500 μm) of solid free form fabrication polypropylene based scaffolds. The results showed optimum proliferation of pre-osteoblastic cell line MC3T3-E1 for scaffolds with 200 and 350 μm pore sizes, whereas the 500 μm pore size scaffolds barely contained cells after 7 days [57]. This suggested that 500 μm pore size was too big for cells to interact with the scaffold under static seeding conditions. In other words, 500 μm pore size enhanced cellular infiltration instead of promoting cell adhesion to the scaffold, as proved for scaffolds with pore size of 350 μm [58]. Similarly, collagen GAG scaffolds with pores ranging from 85 to 325 μm showed the highest osteoblast adhesion for the largest pore size (i.e. 325 μm) [19]. Nevertheless, an earlier work showed that a smaller pore size was better for osteoblast initial adhesion, showing that pores of 95 μm were better for initial attachment than pores bigger than 150 μm [59]. Opposed to these results, coralline scaffolds with two different pore sizes showed that 500 μm pore size presented higher rates of proliferation compared to 200 μm pore size, although the differentiation was

Fig. 3. a) Model and SEM images for heterogeneous (Grad) and homogeneous (Homog) scaffolds produced by 3D plotting techniques, b) stress vs. strain, c) Young's modulus, and d) % of recovery (image from Sobral et al. [50]).

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

higher for the 200 μm [60]. The differences in the results with the previous works may be ascribed to the fact that the coralline scaffolds are ceramic based and therefore controlling the interconnections size is more complex than in the polymeric based scaffolds. Therefore, as the pore size increased up to 500 μm, the interconnection became sufficiently big for cells to penetrate, whereas when the pore size was 200 μm, cells were not able to penetrate and hence showing lower values of proliferation. Since the proliferation was restricted, cells were forced to commit their fate into a differentiated morphology. The pore size has also been shown to have an effect on the proliferation and differentiation of cells for cartilage tissue engineering. Adipose stem cells were seeded on PCL scaffolds prepared with different pore sizes (100, 200 and 400 μm) and were placed under chondrogenic differentiation conditions for 21 days. The results showed that proliferation was higher for the 100 and 200 μm pore sizes, whereas cells tended to agglomerate in the 400 μm pore size scaffolds. Nevertheless, proteoglycan production as well as chondrogenic markers were significantly higher for the 400 μm pore size scaffolds compared to the 100 and 200 μm pore sizes [61]. Similarly, chitosan-hyaluronic acid scaffold showed that chondrocytes expressed higher amounts of collagen II and GAG in the bigger pore sizes (400 μm) compared to the smaller ones (100 μm) [62]. On the same direction, when chondrocytes were cultured in gelatin scaffolds, the results showed that the smaller pore sizes (50–150 μm) caused higher cell number at 21 days, whereas the bigger pore size (350–500 μm) allowed higher production of GAGs. These observations are coupled with the fact that the bigger pore size permitted normal chondrocytes function and normal proliferation up to 14 days, followed by a decrease in the proliferation and an increase

927

in the GAG production [36]. Another work seeded chondrocytes in chitosan scaffolds with different pore size ranges, being the sizes smaller than 10 μm (Fig. 4a), between 10–50 μm (Fig. 4b) and between 70–120 μm (Fig. 4c), which were then transferred into a bioreactor [63]. After 4 weeks, the scaffolds with larger pore sizes (70–120 μm) (Fig. 4f) contained a richer production of collagen II and GAG compared to the smallest pores (10 μm) (Fig. 4d) and the 10–50 μm pores (Fig. 4e). This fact was attributed to an improved diffusion of both nutrients and cells [63]. Cell aggregation and cell–cell contact is known to be the most significant step for chondrogenic differentiation. Hence, the higher pore size allow the allocation of higher number of cells in the pores, which tend to agglomerate once they encounter other cells in the bigger pores showing higher markers of chondrogenesis. In the smaller pores, the number of cells is more limited and therefore preferentially attach to the substrates rather than surrounding cells since there is no proper space to accommodate more cells. Another effective system to determine the optimum pore size is through the use of gradient scaffolds. In general, results show a cell dependent behavior, presenting osteoblast and chondrocytes preferentially in the bigger pore sizes after 56 days, whereas fibroblasts are mainly present in the smaller ones [64,65]. Furthermore, adipose stem cells grown in the gradient pore size (90–400 μm) scaffolds were shown to have the highest chondrogenic differentiation but lowest proliferation in the biggest pore size (400 μm) [66]. This further confirms the previous hypothesis, showing that when cells encounter bigger pore sizes, enough number of cells can be allocated to allow the cell–cell contact and hence present higher markers of chondrogenic differentiation.

Fig. 4. a–c: SEM images and d–f: histology (Hematoxylin and eosin stain, ×40) of chitosan scaffolds 4 weeks after a dynamic culture. The pores diameter of the scaffold were a, d) b10 μm, b, e) 10–50 μm, and c, f) 70–120 μm (image adapted from Griffon et al. [63]).

928

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

3.2.2. Interconnections and permeability Interconnections refer to the pore size that connects two different pores which is directly related with the permeability, which is considered as the ability to allow the circulation of a fluid (e.g. nutrients) in the scaffold. As an illustrative example, rapid prototyping scaffolds with well-defined interconnected pores have been shown to allow nutrient inflow and metabolic waste outflow in a much more efficient manner compared to salt-leached scaffolds, proving to be a better scaffold for cell ingrowth [67]. In order to purely study the effect of the interconnections on the overall cell behavior, calcium phosphates based scaffolds with interconnections ranging from 30 to 100 μm were studied. When osteoblasts were cultured for 14 and 28 days, the cells could penetrate within the different interconnections studied allowing the full colonization of the pores, 40 μm being the most favorable interconnection size [68]. Griffon et al. evaluated the effect of interconnective pore size for the proliferation of chondrocyte using chitosan sponges, showing that chondrocyte proliferation and metabolic activity increased as the pore size interconnection was increased [63]. In this sense, the permeability can be directly correlated with the interconnectivity of the pores, although the pore morphology may also have direct implications in the permeability of the scaffolds. In this sense, a 3D-printed gyroid pore structure was compared with the random pore architecture from salt leaching scaffolds, showing a 10-fold higher permeability for the gyroid pore scaffolds, which could explain the higher number of cells growing in the center of the scaffolds. Opposite to this behavior, the salt leached scaffolds presented a cell sheet on the surface with no cells in the center [69]. Permeability also plays a key role in cartilage tissue engineering. For PCL scaffolds fabricated with different permeability obtained by a tailormade processing, chondrocytes produced more cartilage matrix when the permeability was reduced. In contrast, when the permeability was higher, there was an increase on chondrogenic differentiation of BMSCs [70]. This can be related with the native structure of cartilage that present rather low levels of vascularization and hence depriving permeability of both nutrient and oxygen. These results suggested that mimicking these low levels of permeability may favor cartilage regeneration.

3.2.3. Pore morphology The study of the pore morphology of scaffolds intends revealing the effect of the pore shape regardless of its size. Poly(ethylene glycol) scaffolds were made by controlling the freezing method to allow the formation of spongy-like scaffolds (random spherical pores) and columnar type scaffolds (vertical channel-like pores). Although cells proliferated at a similar rate in both type of scaffolds, higher levels of osteogenic markers were found in the spongy-like scaffolds, indicating that random spherical pores enhanced osteogenic differentiation [71]. When comparing columnar pores with lamellar ones prepared by unidirectional freezing, columnar scaffolds with larger pore width provided higher cell proliferation and differentiation [72]. In a similar work for tendon reconstruction, collagen-GAG scaffolds were prepared by a directional solidification. The results showed that when tendon cells were cultured on these scaffolds, higher levels of attachment, metabolic activity and cellular alignment were observed for those scaffolds with directional pores. It is expected that the higher mobility of cells in these types of scaffolds, as well as higher space for cells to spread and proliferate while maintaining a homogenous supply of nutrients, may enhance the cellular behavior. Opposed to this, the conventional scaffolds with rather smaller interconnections may limit the surface available for cells to migrate and proliferate, leading to cellular aggregation and consequently presenting a different behavior [73]. The development of rapid prototyping scaffolds, which pores shape can be capriciously designed, can provide important information regarding pore morphology [69]. The effect of the pore morphology of 3D-printed structures on the final cell proliferation and differentiation was studied for four different scaffold morphologies: basic (Fig. 5a), basic offset (Fig. 5b), crossed (Fig. 5c) and crossed offset (Fig. 5d). The results showed that the proliferation was higher on the offset scaffolds, whereas differentiation was higher on the basic scaffolds [74]. The presence of rather homogenous scaffolds (basic scaffolds) allows cells to homogenously populate the scaffold, being able to penetrate as cell growth takes place. On the offset scaffolds, cells have a different sensing mechanism, since the offset struts offer more anchorage points for cells than the basic scaffolds. Besides the conformation of the struts gives higher anchorage points, the angles in which the struts interconnect

Fig. 5. SEM images (bars represent 2 mm) of PCL scaffolds produced by a 3D plotting technique with different architectures: a) basic, b) basic-offset, c) crossed and d) crossed-offset, including μ-CT images (inset pictures) (image from Yilgor et al. [74]).

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

are different to those in the basic scaffolds, which may also be a factor that influences differentiation, since it is known that pore morphology and shape is able to alter the cellular behavior [70]. The morphology of the pores has shown to modify the cell behavior for cartilage regeneration as well. Fibroblasts, muscle cells and epithelial cells were cultured in scaffolds fabricated with poly(1,8-Octanediol-cocitrate) polymer with spherical and cubical pores. The results showed that spherical pores with low permeability had enhanced chondrogenic performance in terms of matrix production and gene expression in vitro compared to the permeable and cubical pores [64]. It is expected that the lower permeability, as previously discussed, may mimic more the structure of cartilage and hence enhance a more chondrogenic behavior. 3.2.4. Pore size distribution The study of the pore size distribution intends understating how cells behave in the presence of significantly different pore sizes in the same scaffold as opposed to scaffolds that present merely one pore size. In natural and synthetic polymers, temperature can control the pore size as well as the pore size distribution [75]. Previous work revealed differences in cell seeding when comparing PCL scaffolds with a monomodal and a bimodal pore size distribution, presenting higher initial adhesion for the bimodal pore size distribution [76]. Despite of the enhanced initial cell seeding efficiency in the bimodal pore distribution, the proliferation and differentiation of MSC at 14 days was higher on the monomodal scaffolds [76]. These observations were related with the spatial distribution of the cells within the scaffold. In the monomodal pore size distribution scaffolds, cells mainly remained on the surface of the scaffold and hence cells were directly exposed to the nutrient supply, allowing them to proliferate faster than in the bimodal pore size distribution scaffolds. Similarly, in another work, scaffolds of PLGA with heterogeneous pore size distributions in the form of pore gradient (very small pores on one end and much bigger pores on the other end), showed an increased cell seeding efficiency in comparison to homogenous ones [50]. The cell increase in the heterogeneous scaffolds was ascribed to the higher chances of the cells to encounter struts. In contrast, the homogenous pore structure allowed for cells to directly flow through the scaffold, having fewer chances of interacting with the scaffolds [50]. In contrast Choi et al. claimed that a homogenous pore size distribution scaffold with high interconnectivity provided a better environment for the cells. The homogeneity of the pores avoided cell concentration on the perimeter of the scaffold, allowing the cells to penetrate and obtain cell nutrients [77]. Opposite to these results, a work revealed that the pore size distribution did not affect cell proliferation. PCL scaffolds made by rapid prototyping having staggered, lattice or and diagonal strut morphologies did not show any differences in the number of cells at 1 or 7 days [45,78]. Gradients scaffolds have also been considered an interesting approach to reconstruct the rather anisotropic structure of cartilage tissue. For this purpose, scaffolds having homogenously spaced pores or pore size gradients can be studied showing that the latter allows the formation of the desired anisotropic cell distribution present in cartilage [79]. 3.3. Preclinical assays (in vivo) The most important outcome of macroporosity is its effect on tissue regeneration. Once implanted, surrounding tissues interact with the biomaterial and allow tissue ingrowth. This integration should be smooth and allow tissue regeneration to grow in a similar way to that of native tissue. In order to enhance this process, the macroarchitecture of the desired tissue needs to be resembled. For instance, for bone tissue regeneration, proper bone ingrowth and blood vessel formation does not take place unless the diameter of the pores are between 100 and 400 μm [4,61,80–82]. Nevertheless, there are other important issues that will be discussed in this section to understand whether homogenous or heterogeneous pore size is more favorable for tissue

929

regeneration, or whether tissue beneficiates from specifically oriented pore morphology versus a more random morphology in in vivo implantations. In order to quantify, in the case of bone, the ability to form new tissue, one of the most powerful non-destructive imaging techniques that allow following the performance of a porous implant in situ is micro computer tomography (μCT). μCT is particularly useful when working with biopsies of small sizes (few millimeters) or small animals [83]. The accuracy of the technique and the multiple analysis options allows a wide range of applications in tissue engineering. For instance, μCT has not only been an impressive tool for in vivo imaging, but also allows to characterize the porosity and pore size distribution of scaffolds [84] or even to quantify the amount of new apatite formation on scaffolds after immersion in simulated body fluid [85]. 3.3.1. Pore size and pore size distribution The pore size range plays a critical role in determining the outcome of tissue regeneration. Several studies have attempted to find out the optimum pore size for in vivo bone regeneration by implanting scaffolds with different pore sizes. For instance, bone regeneration has been achieved with nickel titanium alloys prepared with custom made pore sizes (259, 272 and 505 μm) and implanted in the distal femoral metaphysis of rats for 30 weeks. The results showed that the optimum pore sizes for bone regeneration were 259 and 505 μm, which presented similar bone regeneration, although the biggest pore sizes (505 μm) showed lower presence of fibrotic tissue [86]. In a similar way, another work reported that pore sizes of 565 μm presented a better bone ingrowth than pores in the range of 300 μm [87]. Since the proper range is yet to be clarified, PLGA scaffolds were used to repair the tibial head of rats with three different pore sizes: 100–300, 300–500 and 500–710 μm for 2 and 4 weeks. The results showed higher novo bone formation for the pore sizes between 300 and 500 μm [80]. In many occasions, the difference in the pore size distribution makes difficult obtaining accurate data of the exact pore size that beneficiates the proper bone regeneration. For this purpose, gradient pore scaffolds are an attractive type of scaffolds that may allow screening a wide range of pore sizes. It was observed that the pore range between 290 and 310 μm showed faster new bone formation compared to smaller and larger pore sizes [65]. As previously mentioned, 3D printing technology serves as a powerful tool to prepare these types of scaffolds with custom made pore sizes. The findings showed that the implantation of the custom made scaffolds in rabbits for up to 16 weeks allowed optimum bone regeneration in the pores which closely resembled the size of trabeculae, which range from 100 to 250 μm [88]. Opposite to these results, an earlier work reported that a pore size of 100 μm is not the threshold value to be able to grow new bone into the scaffold [89]. In this sense, Ti shaped plates with different pore sizes, ranging from 50 to 125 μm, were studied in the cancellous bone of the distal rabbit femur for 12 weeks. The results showed that the amount of new ingrown bone was independent of the pore size, showing osteonal structures even in the smallest pores [89]. In accordance to these results, another work found no differences in the in vivo bone formation response for β-TCP scaffolds with four different pore sizes (150, 260, 510 and 1220 μm). Nevertheless, a faster resorption was observed for the scaffolds with bigger pore sizes, although it was associated with lower bone content and higher soft tissue content [90]. Similar results showed that PCL scaffolds made by rapid prototyping with 3 different sizes (350, 550 and 800 μm) had few differences in bone regeneration. This encourages the future studies of other scaffold structural properties such as the pore shape, interconnectivity or permeability [91]. Taking into account that bone regeneration may be seen as a two stage process which involves initial angiogenesis followed by a later osteogenesis, it has also recently become of great importance to focus the design of scaffolds to enhance the initial angiogenesis [92,93]. Angiogenesis is known to be a key step for the nutrient and oxygen transport that may enhance the overall regeneration process. In order to see the

930

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

effect of the pore size on angiogenesis, calcium phosphate particles with different pore sizes (40–70 μm, 70–140 μm, 140–210 μm and 210– 280 μm) were implanted for 28 days in critical size cranial defects. The results showed that the pores that were bigger than 140–280 μm had higher capillary density than the smaller pore sizes (40–140 μm). The volume of new bone deposited was pore size dependent, having the highest value for the 210–280 μm pore size range. The authors hypothesize that angiogenesis is essential for the subsequent osteogenesis and hence, big pore sizes are needed [81]. Despite having the proper porosity is essential for angiogenesis to take place, chemical ions may synergistically enhance the angiogenic cascade to take place. Among the ions that are known to enhance angiogenesis are those that mimic hypoxic conditions, mainly cobalt and copper [1,94,95]. The different ions can easily be incorporated within the structure of the scaffolds and be released at initial times, which needs to be carefully designed to avoid high ion concentration microenvironments that are known to increase cytotoxicity [93]. While the angiogenic initiators arise from a chemical stimulation and play a key role in angiogenesis, the proper guidance of the new blood vessels takes place through the architecture of the porosity. Therefore, a synergistic effect is obtained when combining the proper signaling with the proper spacing for cells to populate the scaffolds. Related with the initial angiogenesis and its effect on bone regeneration, β-TCP scaffolds were prepared with different pore sizes but with the same interconnections. The scaffolds were prepared with pore sizes between 300–400 μm (Fig. 6a), 400–500 μm (Fig. 6b), 500–600 μm (Fig. 6c) and 600–700 μm (Fig. 6d). The results showed that the amount of fibrous tissue created decreased by increasing the pore size. It was hypothesized that bigger pore sizes were important for proper blood vessel formation, since pores smaller than 400 μm (Fig. 6a) could considerably limit blood vessel formation, allowing only smaller vessels, whereas bigger pore sizes allowed proper new blood vessel formation (Fig. 6b–d) [96]. Besides bone, osseochondral defects also need for the proper scaffolds to be placed. Nevertheless, the different nature of the cartilage and bone demands different pore sizes. In this sense, bilayered scaffolds with different pore sizes are used to simultaneously heal chondral and osseous defects. It is interesting to observe that pores in the range of

100–200 μm are optimum for the chondral repair, whereas pores in the range of 300–450 μm are optimum for the osseous layer [82]. Similarly, PCL scaffolds with different sizes ranging from 100 to 400 μm were investigated for cartilage repair and the results showed that the 200 and 400 μm pore size gave better results than the 100 μm [61]. 3.3.2. Interconnections and permeability Scaffolds architecture may also be designed in a way that their permeability is maximized. In general, better bone penetration and vessel infiltration can take place when scaffolds present higher permeability, suggesting that higher permeability with regular architecture is better for bone regeneration [97,98]. The permeability is closely related to the pores interconnection. Interconnectivity has been shown to play a dominant role in the ability of fluid and cells to penetrate and hence low interconnectivity may impair tissue repair [99]. In order to proof the importance of the interconnections for bone regeneration, a calcium phosphate scaffold with different interconnections was implanted in the middle shaft of rabbit femurs for 12 and 24 weeks. Pore interconnection of 20 μm was shown to be enough for cells to penetrate, although the size had to be over 50 μm to allow new bone ingrowth [68]. 3.3.3. Pore morphology Pore morphology is of great importance for bone regeneration since it relates with the ability of the scaffold to integrate within the tissue. The way in which pores are allocated in reference to the tissue and the ability of the scaffold to mimic the structure of bone may be some of the parameters that can lead to a better design of scaffolds. Higher cell infiltration and vascularization is observed in columnar scaffolds compared to conventional sponge-like scaffolds, hereby proving the importance of the pore morphology [71]. In a similar way, when implanting collagen scaffolds into a surgically created rat diaphragm defect, the scaffolds with radial pores (Fig. 7a, b), obtained by inward out freezing, are better integrated within the surrounding tissue than the round random pore scaffolds (Fig. 7e, f), made with conventional freeze-drying techniques. Furthermore, cells were able to infiltrate further into the center of the radial scaffolds, allowing for blood vessels to migrate deeper inside the scaffolds [100]. Similarly, collagen radial

Fig. 6. Light microscope images obtained 4 weeks after implantation of scaffolds with different pore sizes: a) 300–400, b) 400–500, c) 500–600 and d) 600–700 μm, with a constant interconnection size of 120 μm. The blue arrows indicate blood vessels. TCP = β-tricalcium phosphate; FT = fibrous tissue (image from Feng et al. [96]).

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

931

Fig. 7. SEM images of radial (a, b) and control (e, f) scaffolds. Microscopic pictures of radial (c, d) and control (g, h) scaffolds stained with hematoxylin and eosin stain. In a and e, the lines indicate the interfaces between the compressed and radial/porous scaffolds parts. In microscopic images, a wavy radial pore orientation was visible in radial scaffolds (d), whereas an overall random round pore structure was visible in control scaffolds (h). Immune cell infiltration (indicated by asterisks) was present at the interface between the scaffold and diaphragm (c, g). The upper side of the pictures is the pulmonary side. di: original diaphragm; li: liver; sc: scaffold. Bars represent: a, c–e, g, h: 200 μm; b, f: 100 μm (image adapted from Brouwer [100]).

oriented scaffolds were fabricated by a temperature gradient-guided thermal-induced phase-separation technique. The scaffolds were implanted in the patellar groove of New Zealand white rabbits to evaluate cartilage regeneration. After 6 and 12 weeks neo-cartilage partly aligned with the direction of the radial scaffold, whereas in the control scaffolds with random pore orientation, bone grew in a random manner. At a closer examination, the oriented pore scaffolds formed a mixture of fibrous and fibrocartilage, whereas the random scaffolds mainly formed fibrous tissue. These results pointed out that the radially oriented structures enhanced a better healing of osteochondral defects [31]. Opposite to these results, implantation of polyurethane scaffolds in rats presenting random oriented pores and channel like pores showed that full tissue ingrowth was achieved only after 8 weeks for the randomly oriented pore scaffolds, whereas for the channel like scaffolds this was not achieved until week 24. Nevertheless, the random pores induced tissue infiltration with unorganized collagen deposition, whereas the channel like pores induced tissue ingrowth and collagen alignment in the direction of the scaffold channels [101]. The faster tissue penetration in the radial pore morphology tends to show regeneration in the center of the implant, whereas a more orthogonal design of the scaffold leads to an interpenetrating matrix of the scaffolds and new bone [102]. In the case of cartilage regeneration, the anisotropy of pores has not been shown to significantly affect its regeneration [103]. In a similar way, round pores can be compared with elongated pores using ceramic phosphate glasses. Its implantation in rabbits showed that round pores enhanced new bone formation, whereas the elongated pores caused a lower resorption. Nevertheless, the results need to be observed with care since the properties of the scaffold in terms of total porosity as well as pore size were significantly different, making these results difficult to interpret [104]. Liu et al. used unidirectional freezing of suspensions to create bioactive glass orientated scaffolds (columnar pores) with a total porosity of 50 vol.% and a pore size diameter 50–150 μm. As controls, trabecular scaffolds with larger porosity (total porosity of 80 vol.%) and larger pore size diameter (100–500 μm) were used. The scaffolds were implanted in bone defects created in the parietal bone of rats and evaluated at 12 and 24 weeks post-surgery. At 24 weeks, a higher amount of new bone (normalized to the available pore volume) was formed in oriented scaffolds (55 ± 5%) compared to trabecular scaffolds (46 ± 13%) despite the bigger pore size of the trabecular scaffolds [105]. This proves that the oriented structure supported better bone ingrowth, not only due to a guiding effect of the pore morphology, but also to a more

favorable physiological environment in terms of metallic ions, proteins and other biological molecules. 3.4. Macroporosity overview The first part of this review has focused on the characteristics of the macropores (i.e. size, interconnectivity, distribution and morphology) that influence the physico-chemical and biological properties of a

Fig. 8. Scheme displaying how structural properties of the scaffold (i.e. interconnectivity, pore size, pore size distribution, and pore and morphology) have an effect on certain physico-chemical (mechanical properties, permeability) and biological (cell seeding/ penetration, in vivo response, cell proliferation and cell differentiation) behaviors. The black internal arrows point out the increase of the structural property. The colored lines correlate the structural property of the scaffold to physico-chemical or biological behaviors; whereas the arrows point out a trend, the dashed lines indicate controversy or a not determined trend.

932

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

scaffold. Fig. 8 summarizes the general mechanical and biological properties of scaffolds and how their pore size, pore size distribution, interconnectivity and pore morphology affect all these parameters. The general trend of each property is given by the direction of the arrows. For instance, the mechanical properties (green color) tend to increase as the pore size is decreased as shown in the figure. In a similar way, as the interconnectivity of the scaffolds is decreased, the mechanical properties tend to increase, whereas spherical pores have as well higher mechanical properties than rather elongated pores. In similar line, monomodal pores size distributions have higher mechanical properties than bimodal pore size distribution. Some other properties have less dramatic tendencies, such as for instance the permeability (dark blue color). It can be observed that as the interconnectivity increases, the permeability is increased as indicated by the arrows. Nevertheless, no direct effect is given when the pore size, pore size distribution or pore morphology are varied as shown with the dotted lines. The biological response of the different types of scaffolds is also influenced by their pore morphology and structure. In this way, the seeding efficiency (red color) was shown to be better in bigger pore sizes, whereas intermediates values of interconnectivity and bimodal homogenous pore size distribution had highest values. Furthermore, radial pore morphology is detrimental for cell seeding, whereas round pore morphology decrease as well as cell seeding efficiency. Despite these differences, their proliferation (light blue color) is not significantly affected by the different parameters, in general not presenting a clear tendency. Nevertheless, cell differentiation (purple color) was shown to be higher for bigger pore sizes with higher interconnectivities. The pore size distribution and pore morphology had less dramatical effects. Finally, the in vivo response (yellow color) is greatly determined by the interconnectivity of the pores, which in turn relates with the permeability, presenting optimum bone formation for interconnected pores. Furthermore, an intermediate pore size of bimodally distributed homogenous pores seems to play key role in regeneration. Despite the pore morphology does not have a significant effect, radial pores seem to slightly increase the bone formation ability. Even though it is difficult to define the exact porosity properties for an ideal scaffold, Hing indicated as a broad guideline that a total porosity should be higher than 50–60 vol.%, an interconnection size should be higher than 50–100 μm and strut porosity should be higher than 20 vol.% [106]. 4. Microporosity The role of macroporosity has been mainly associated with the ability of a scaffold to allow proper bone ingrowth and bone regeneration, while having in general slight effects on cell proliferation and differentiation. Nevertheless, the porosity in the range of nanometers up to several

microns has tremendous effects on the ability of cells to proliferate and differentiate, and hence play a key role in the overall bone regeneration. Not only it is able to regulate the phenotype of cells to induce higher cell mineralization, but also increase the protein adsorption, which in turn can increase the osteoinductive capacity of a material. Microporosity can be incorporated into ceramic and polymeric scaffolds with different techniques. Some of these are indicated and briefly described on Table 2. 4.1. Tailoring osteoinductivity through microporosity — the importance of protein interaction with microporous substrates In the last decades, the interest for biomaterials that are not only able to support cell colonization and thus regeneration of tissue, but that are also able to induce bone formation even when implanted at heterotopic sites has exponentially grown. These materials have been widely studied in the bone regeneration field and are known as osteoinductive biomaterials [14]. The reason why certain biomaterials are osteoinductive whereas other chemically identical materials are not has been the subject of a large number of scientific discussions. In early 1990s, certain hydroxyapatite ceramics [116] as well as other porous calcium phosphate ceramics [117,118] were reported to form bone tissue when implanted subcutaneously. Yamasaki et al. highlighted that the presence of interconnected micropores (2–10 μm) were able to confer the scaffolds with osteoinductive properties, since bone ingrowth was not observed in analogous materials with dense morphology [118]. The ability of a scaffold implanted into the body to enhance bone formation is associated to the entrapment and adsorption of certain type of proteins circulating around the biomaterial. Cells are able to recognize specific peptides domains found within this proteins to initially attach and to control their cell fate at a later state [119]. This recognition mechanism takes place through a family of protein receptors called integrins, and their main role is to transmit information across the cell membrane by linking the macromolecules of the extracellular matrix with the cytoskeleton of the cell [120,121]. Based on the importance of microporosity to entrap and adsorb proteins, and hence to enhance bone formation, several studies have closely examined this phenomenon. Wei et al. [122] developed fibrous scaffolds of PLLA/PDLLA and PLLA/PLGA blends with tailored porosity and roughness. The results showed that the capacity to adsorb proteins was enhanced when the total specific surface area of the scaffold was increased. A similar study also evaluated the ability of ceramics with different chemistry (HA and BCP) and porosity to adsorb proteins. The capacity to adsorb fibronectin and vitronectin was evaluated with materials with pores around 200–1000 nm [123], whereas materials with pores around 20–80 nm were used to evaluate the adsorption of

Table 2 Methods to incorporate micropores into ceramic and polymeric scaffolds. Material

Method

Description method

Refs

Ceramic Ceramic Ceramic Ceramic cements

Polishing Sinterization of slurries with organic compounds Sintering of powder with pressure Modulation of liquid-to-powder ratio

[107] [108] [109] [110]

Ceramic, polymer

Freeze drying

Ceramic, polymer

Carbonate in matrix

Polymer

Plasma etching (superficial)

Polymer Polymer

Thin film deposition Casting

Polymer

Cast and freeze-extraction method

Surface is polished with a rough paper Synthesis of ceramic slurries followed by filtering, drying and sintering Powder pressed uniaxial at different sintering conditions (temperature and time) When crystals precipitate voids are left between them, the amount of porosity being proportional to the water present in the paste An aqueous matrix is frozen and afterwards dried by sublimating the ice under vacuum in a freeze-dryer Organic matrix containing different amounts of carbonate salt (e.g. ammonium bicarbonate). Plasma (partially ionized gases) is applied to the surface of a material modifying its structure. Plasma is used to obtain thin hard films such as carbon nanotubes, ZnO or TiO2. Polymer in an organic solvent is casted on a silicon wafer surface and kept under vacuum for several days Microbead-template is filled with a polymer solution. Frozen and removed by solvent exchange

[111] [112] [113] [113] [114] [115]

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

fibrinogen and insulin [108]. In both studies, the higher amount of micropores as well as the higher specific surface area of the materials caused an increase in the adsorption of the serum proteins. However, another protein found in serum with larger dimensions (e.g. collagen type I), was less sensitive to the different microporosities (20–80 μm pores) [108]. These results prove that the microporosity of a biomaterial may be as important as the competitive effects described by the “Vroman effect” to determine which serum proteins adsorb onto a biomaterial, which will in turn determine cell adhesion. A family of interesting materials that are based on a cementation reaction and that can lead to significant different micron pore sizes depending on the initial size of the particles are the calcium phosphate cements [110,124–126]. Similar studies were performed to analyze the protein adsorption in calcium deficient hydroxyapatite substrates with a broad pore size distribution at the micron (1–5 μm) and sub-micron range (b 1 μm) [110]. The authors revealed that the micron pore size distribution allowed higher protein penetration and adsorption compared to the sub-micron substrates. The results rely on the fact that the sub-micron pores presented a lower pore interconnection and hence limited the protein penetration. Moreover, the higher pore size distribution allowed improved protein penetration by avoiding strong electrostatic interactions with the pores [110]. These substrates with significantly different topography can also serve to evaluate the entrapment and desorption of proteins through size exclusion chromatography [127]. The authors concluded that the substrates composed by a low packing density of large CDHA crystals allowed more protein concentration and a faster elution time in comparison with a configuration of highly packed small CDHA needles. Besides of the biological molecules, inorganic ions released by a biomaterial to the physiological environment can also have a crucial effect on determining its osteoinductive properties. Some of these ions (silicon, calcium, phosphorus, zinc, magnesium, strontium, copper, among others) may cause the recruitment and regulation of cells around the biomaterial [106,94]. This is the reason why bioactive glass ceramics are commonly considered to have osteoinductive properties (also known as osteoproductive properties) [128]. For more insight about the cell behavior triggered by different metallic ions, the readers are addressed to the excellent reviews written by Hoppe et al. and Mourino et al. [94,95]. In summary, the presence of micron sized pores in the surface may allow the adsorption of sufficient levels of specific proteins (e.g. bone morphogenic proteins, BMPs) that can potentially confer osteoinductive properties to the matrixes [15,129–131]. Fig. 9 schematically shows

933

how the different microporosities on the surface are able to bind different amount of serum proteins, which in turn translate into different levels of cell interactions. The higher roughness allows for an enhanced cell spreading which is related with the higher focal adhesion sites for cells to attach and interact with the material. Moreover, the orientation and conformation of proteins adsorbed onto surfaces is critical for cells to recognize specific active domains on the adsorbed proteins, which may initiate signaling events. These biomolecular cues can determine cellular behaviors such as adhesion, morphology, migration, proliferation and differentiation [132]. It is important to highlight that the accessibility of the proteins for the surface will depend on topography and size impediments. Only those proteins whose size allows them reaching the surface of a biomaterial will be able to adsorb onto it and in turn affect those cells whose size will also allow reaching that area. 4.2. Micropores in drug delivery As mentioned earlier, micropores play a key role in controlling protein adsorption as well as cell–material interactions. Nevertheless, these pores may also be efficient systems for the loading and release of specific biological molecules with regenerative potential. While scaffolds can be designed to present certain macroporosity, depending on the processing parameters scaffolds can tune their porosity also at the micron level. These pores which are usually in the range of tens microns have been shown to be ideal for growth factor allocation. Depending on the processing parameters, the molecules can be incorporated inside the scaffold or loaded afterwards through an adsorption method. Since several polymeric scaffolds rely on the use of organic solvents for their preparation while ceramics usually use high temperature for the stabilization of the scaffold, a common strategy to load the molecules is by adsorption. As exemplar studies, ceramic bioglass based materials possess a unique microporous structure that is intrinsically present in these types of materials. This has been shown to have unique loading and very sustained release of biological molecules. Generally these scaffolds need to be prepared at high temperature in the presence of a template and porogens that are removed at high temperature to obtain the micro (~ 6 nm) and macroporosity (200–500 μm). These in general have shown excellent loading of drug that have enhanced bone regeneration [133]. The absence of the templates and porogens that lead to micropore formation in these scaffolds has been shown to restrict growth factor loading and hence perform in a very limited manner, due to lower pore volume and surface area, showing pores as small as 4 nm in the absence of

Fig. 9. Scheme correlating microporosity with cell behavior. A higher of microporosity in a biomaterial favors the absorption of a large number of proteins, which in turn causes specific interactions of cells.

934

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

porogens which may be increased up to 10 nm with 10 fold increase in the pore area when porogens are present [134,135]. Other sources for ceramics scaffold fabrication are the calcium phosphate based scaffolds, which are easily prepared by combining hydroxyapatite with a polymeric material, e.g. PMMA, which is burned after the processing to obtain micropores in the range of several microns [136,137]. The presence of the micropores (~5 μm) allowed enhanced BMP-2 adsorption and hence enhances bone regeneration [137]. In a similar way, the intrinsic microporosity of precipitated HA is also known to be excellent source for high adsorption and controlled release of growth factors [126]. Other studies used the micropores in polymeric treated scaffolds to allocate the drugs. For instance, PLA scaffolds can be plotted into 3D scaffolds with controlled porosities designed by computer. The PLA biopolymer solution can be combined with temperature ionic liquid incorporating certain ions that once the scaffold is plotted and immersed in ethanol, the ionic liquid is completely removed to allow formation of micropores (~2.5 μm). Due to the presence of microporosity, the loading of model molecules was shown to be up to 10 times higher compared to the non-microporous scaffold, which then showed a sustained release over a month for the microporous scaffolds [138].

While most of the described ceramic bioglasses need high temperature processes for the mechanical stability, the use of setting mechanism to increase the mechanical properties is an interesting system to obtain encapsulated molecules that can be delivered through their micropores. For instance, silica based materials that are prepared by sol–gel reaction offer a nice system to incorporate growth factors in mild conditions that can then be released in an almost zero order kinetics (pore sizes ~ 10 nm) [139–141]. In a similar way, another type of materials that allow a setting reaction are the calcium phosphate cements. As previously described, these present an intrinsic micron sized porosity (1–5 μm) and due to the mild conditions allows the incorporation of factors during their preparation [125,142,143]. In general, their release profile has been shown to be sustained, although it has been previously shown that the formation of hydroxyapatite during the cementation reaction may block some pores, trapping the molecules inside these [144–146]. 4.3. Cell behavior on microporous materials Surface topography and microporosity of a material also plays a very important role on cell behavior. The influence of pore sizes on cell

Fig. 10. Morphology of MG-63 osteoblast-like cells cultured for 1 day on polycarbonate membrane surfaces, whose diameter pore sizes were: a) 0.2, b) 0.4, c) 1.0, d) 3.0, e) 5.0, and f) 8.0 mm (original magnification ×3000) (image from Lee et al. [57]).

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

behavior was determined by culturing MG-63 osteoblast-like cells on polycarbonate membranes designed with different pore sizes. The cells spread and adhered better on membranes with smaller micropores (0.2 μm diameter) (Fig. 10a, c and e) than on those with larger micropores (3.0–8.0 μm) (Fig. 10b, d and f). Moreover, the cells cultured on larger micropores produced increased levels of ALP and osteocalcin [147]. In another study, the different microporosity and topography of CDHA materials (total porosity of 35 vol.%, pores b 5 μm) was partially responsible for the different patterns of proliferation and differentiation observed for osteoblast cells. Materials with smaller CDHA crystals stimulated differentiation, whereas those with bigger crystals enhanced proliferation [125]. The effect of the amount of pores has been compared in several scaffolds with different chemistry: made of poly ε-caprolactone (PCL) with a total porosity of 80–93 vol.% (pores within 200–300 μm, ~50 μm and b10 μm) [148], made of PLGA 50/50 with a total porosity of 69–81 vol.% (pore size within 5–20 μm) [112], and made of polyethylene terephthalate (porosity of 93–97 vol.%) [149]. The results of these works indicated an improved cell ingrowth in the scaffolds with higher porosity and better interconnected structures. These results were mainly attributed to a higher surface area for cell growth as well to an easier diffusion of oxygen and nutrients, and clearance of waste products [112,149,150]. There are few in vitro studies reporting that low amount of micropores or small pores enhance biological response in front of larger amount of pores or bigger size. Akin et al. proved that proliferation of human bone-derived cells (hBDC) was enhanced in titanium oxide (TiO2) films with smaller diametrical pores (0.5 and 16 μm) in comparison with higher pores (50 μm) [151]. Isaac et al. reported that a higher amount of micropores in scaffolds (0, 25 and 45 vol.%) decreased both the viability and osteoblastic differentiation (ALP activity and osteocalcin secretion decreased) of human bone marrow stromal cells (hBMSC) in comparison with non-microporous ceramics [152]. The authors speculated that the micropores increased the ion exchange between the ceramic and the medium, modulating the calcium and phosphate concentration. Ionic concentration could in turn influence cell spreading, viability and differentiation [152]. In another work, a higher proliferation and differentiation of BMSC was also observed on materials with lower microporosity (5 and 15 vol.% compared to 30 vol.%). In this case, the authors hypothesized that the proteins adhere in low microporosity materials in such orientation that could not stimulate osteoblastic adhesion and proliferation [109]. Regarding the cell differentiation, Kasten et al. [153] observed that ALP activity was independent of the total porosity of the scaffold, whereas Takahashi et al. reported an inverse relationship between these two parameters [149]. In the latter study, the inhibition of cells differentiation was attributed to their enhanced proliferation [149]. Several works have shown that osteoblasts respond in a different manner depending on the superficial roughness of the materials. Wan et al. grew OCT-1 osteoblast-like cells on polystyrene (PS) films with big (2.2 μm) or small (0.45 μm) islands on its surface and showed that cell adhesion was improved on the surfaces with small or big islands in comparison with smooth surfaces [154]. Regarding the cell growth, Anselme et al. observed that a larger roughness (between 0.16 and 3.40 μm) of Ti alloy (Ti6Al4V) decreased cell proliferation of primary human osteoblasts [155] and pre-osteoblasts MC3T3-E1 [156]. Interestingly, Washburn et al. showed similar results for polymeric samples with nanometric roughness. Pre-osteoblasts MC3T3-E1 cells cultured on films of poly(L-lactic) acid proliferated slowly on surfaces with more roughness (13 nm) in comparison with smoother surfaces (0.5 nm) [114]. Opposite results were observed by Deligianni et al. when culturing primary human bone marrow cells on sintered hydroxyapatite disks with roughness between 0.7 and 4.3 μm, since a higher cell adhesion and proliferation was observed on materials with larger roughness [107]. Regarding the differentiation, most works have shown that microscale roughness improves the differentiation of

935

osteoblasts. Osteoblast MG-63 cells cultured on pure titanium and titanium alloy (Ti6Al4V) samples with different superficial roughness (between 0.22 and 3.2 μm) produced more ALP than smooth surfaces [157]. Similarly, Gittens et al. cultured osteoblasts MG-63 on titanium surfaces with different roughness, machined (0.43 μm) or sandblasted/acid-etched (3.29 μm), and showed that MG-63 produced higher amounts of osteocalcin and osteoprotegerin on the rougher samples [158]. Zhao et al. showed similar results by growing MG-63 cells on titanium disks with different roughness: smooth (0.2 μm), acid-etched surfaces (0.83 μm) and sandblasted/acid-etched surfaces (3–4 μm). The larger the roughness of the surfaces, the more osteocalcin was produced by the cells [159]. In contrast, Zhu et al. did not observe any difference on the production of ALP by osteoblasts Saos-2 cultured on titanium samples with different roughness between 0.18 and 0.35 μm [160]. 4.4. Preclinical assays (in vivo) Several experiments have evaluated the effect of the microporosity in in vivo studies. As previously discussed, the increased ability to entrap proteins in the micron-sized pores may stimulate the osteoinductive and osteoconductive properties of the scaffolds. 4.4.1. Osteoinduction (subcutaneous models) Habibovic et al. performed an elegant experiment to determine the role of microporosity in two families of chemically identical porous ceramics: hydroxyapatite (HA) and biphasic calcium phosphate (BCP) [130]. Sintering temperatures between 1100–1200 °C allowed modifying the microporosity (within a pore diameter range b 2 μm) while not altering their macroporosity (249 ± 38 μm). The results showed that the implantation into the back muscles of Dutch milk goats for 6 and 12 weeks allowed bone formation in the presence of micropores but failed when the amount of micropores remained low (HA sintered at 1250 °C). The higher amount of adsorbed/entrapped proteins in the microporous walls enhanced bone formation [130], which was essential to provide the biomaterial with osteoinductive capacity. It is thus speculated that the microporosity modifies the dynamic interface of materials and consequently triggered the differentiation of relevant cells towards the osteogenic lineages [130]. Habibovic et al. also pointed out that a higher microporosity was inherently linked with a higher specific surface area, and hence could cause a major dissolution of ions. The higher ion dissolution would facilitate the apatite formation in vivo, causing the coprecipitation of endogenous proteins (e.g. BMPs) that could in turn trigger the differentiation of recruited undifferentiated cells towards the osteogenic lineage [15,130]. Another hypothesis suggested that the inflammatory response triggered after the implantation of a biomaterial, which causes the release of cytokines that promote the differentiation of circulating mesenchymal stem cells (MSC) into osteoblasts, would induce osteoinductivity [14]. It should be noted that although the osteoinductivity of a biomaterial depends on its physico-chemical and structural properties, the methodology employed to evaluate this property (i.e. animal model, implantation site and duration of the study, among others) also plays a fundamental role [14]. The readers are addressed to an excellent review from Barradas et al. that accurately discusses each of these parameters [14]. 4.4.2. Osteo conduction (bone and osteochondral defects) Ikeda et al. implanted PLGA scaffolds with different proportions of micropores (80, 85 and 92 vol.%) in osteochondral defects made in the femoral condyle of rabbits. After 6 and 12 weeks, scaffolds with higher microporosity (85 and 92 vol.%) replaced a larger volume of the scaffold by bone and cartilage [161]. Moreover, bone marrow cells were also able to migrate into scaffolds with higher porosity (85 and 92 vol.%), which allowed repairing osteochondral defects. Hing et al. implanted porous hydroxyapatite materials in the femoral condyle of rabbits and

936

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

found that while no differences on bone formation were observed between samples with different ranges of macroporosity, a larger volume of new bone with higher bone density were observed in samples with higher microporosity. The authors related these results to the increased permeability, which improved nutrients transfer and waste removal, leading to faster bone apposition [162]. While the aforementioned studies were performed in small animal models, a few studies have introduced microporous implants into big animal models such as dog or human. Story et al. evaluated the formation of bone in microporous implants placed in mandibular and femoral defects created in a dog model [163]. A major bone formation was observed in the implants with the highest porosity (48 vol.%) compared to the less porous ones (44 vol.%) after 2, 4, 8 and 12 weeks of implantation. The authors indicated that the sample with the highest microporosity caused a higher initial bone attachment strength due to an increased bone ingrowth in early healing stages [163]. Knabe et al. introduced a mixture of commercial β-TCP porous particles and autogenous bone chips into humans in a clinical trial. The results confirmed the ones obtained in preclinical trials, showing that higher microporosity in the materials enhanced bone formation and particle degradation [164].

5. Concluding remarks Several parameters have significant importance regarding the porosity in scaffolds, especially focusing in the field of bone regeneration. Not only the macroporosity has been shown to be important for bone regeneration, but so has the microporosity. Macroporosity is the key element to allow cell penetration within the scaffold and hence the formation of new tissue. Therefore, the macroporosity may dictate the extent of regeneration. From our point of view, both the scaffold architecture (morphology and pore size distribution) and the pore size play an important role in the formation of new tissue. Scaffolds with a high number of homogenous pore size allows a faster colonization, as long as these pores are big enough in size. Heterogeneous pore size distribution may also allow cell colonization, although the higher proportion of low pore size may reduce the diffusion of nutrients, oxygen and cellular waste. Actually, controlling the ability of a fluid to flow through the scaffolds, which is regulated by the pore interconnectivity and permeability, is a must in scaffolds to allow proper cell infiltration within them. Cell penetration can be further enhanced by managing the pore morphology, which may promote the homogenous colonization of the scaffold. Furthermore, the development of new processing technologies, such as 3D-printing, allows unique control of the pore morphology, and consequently is a very powerful tool to study thoroughly. Nevertheless, focusing on bone tissue engineering, most of the 3D-printed scaffolds are made of polymeric materials that in general do not possess proper osteoconductive or osteoinductive cues. Therefore, it is of great importance the development of new scaffolds with controlled porosity, which at the same time possess similar chemical and microstructural properties to that of bone, e.g. hydroxyapatite precipitated at body temperature. On the other hand, the microporosity is the one controlling the final cell fate. In this sense, increasing the amount of micropores subsequently increases the interactions with serum proteins, which may be an attractive strategy to promote the osteoinductive and osteoconductive properties of the scaffold. Although the microporosity is often overlooked in scaffolds, it can play pivotal role in obtaining the desired bone functions.

Acknowledgments The present research was conducted with the support of the research fund of Dankook University.

References [1] R.A. Perez, S.-J. Seo, J.-E. Won, E.-J. Lee, J.-H. Jang, J.C. Knowles, et al., Therapeutically relevant aspects in bone repair and regeneration, Mater. Today (2015)http://dx. doi.org/10.1016/j.mattod.2015.06.011. [2] S.J. Hollister, Porous scaffold design for tissue engineering, Nat. Mater. 4 (2005) 518–524, http://dx.doi.org/10.1038/nmat1421. [3] D.W. Hutmacher, Scaffolds in tissue engineering bone and cartilage, Biomaterials 21 (2000) 2529–2543, http://dx.doi.org/10.1016/S0142-9612(00)00121-6. [4] R.A. Pérez, J.-E. Won, J.C. Knowles, H.-W. Kim, Naturally and synthetic smart composite biomaterials for tissue regeneration, Adv. Drug Deliv. Rev. 65 (2013) 471–496, http://dx.doi.org/10.1016/j.addr.2012.03.009. [5] Scaffolding in Tissue Engineering, CRC Press, 2005 (http://books.google.com/ books?id=kSczI9Sbq9EC&pgis=1 (accessed July 17, 2014)). [6] K. Li, X. Li, G.W. Irwin, G. He (Eds.), Life System Modeling and Simulation, Springer Berlin Heidelberg, Berlin, Heidelberg, 2007http://dx.doi.org/10.1007/978-3-54074771-0. [7] K.M. Woo, V.J. Chen, P.X. Ma, Nano-fibrous scaffolding architecture selectively enhances protein adsorption contributing to cell attachment, J. Biomed. Mater. Res. A 67 (2003) 531–537, http://dx.doi.org/10.1002/jbm.a.10098. [8] A. Macchetta, I.G. Turner, C.R. Bowen, Fabrication of HA/TCP scaffolds with a graded and porous structure using a camphene-based freeze-casting method, Acta Biomater. 5 (2009) 1319–1327, http://dx.doi.org/10.1016/j.actbio.2008.11.009. [9] W. Bonfield, Designing porous scaffolds for tissue engineering, Philos. Transact. A Math. Phys. Eng. Sci. 364 (2006) 227–232, http://dx.doi.org/10.1098/ rsta.2005.1692. [10] K. Rechendorff, M.B. Hovgaard, M. Foss, V.P. Zhdanov, F. Besenbacher, Enhancement of protein adsorption induced by surface roughness, Langmuir 22 (2006) 10885–10888, http://dx.doi.org/10.1021/la0621923. [11] J.Y. Martin, Z. Schwartz, T.W. Hummert, D.M. Schraub, J. Simpson, J. Lankford, et al., Effect of titanium surface roughness on proliferation, differentiation, and protein synthesis of human osteoblast-like cells (MG63), J. Biomed. Mater. Res. 29 (1995) 389–401, http://dx.doi.org/10.1002/jbm.820290314. [12] H. Yuan, Z. Yang, Y. Li, X. Zhang, J.D. De Bruijn, K. De Groot, Osteoinduction by calcium phosphate biomaterials, J. Mater. Sci. Mater. Med. 9 (1998) 723–726 (http:// www.ncbi.nlm.nih.gov/pubmed/15348929 (accessed June 5, 2014)). [13] J. Zhang, D. Barbieri, H. ten Hoopen, J.D. de Bruijn, C.A. van Blitterswijk, H. Yuan, Microporous calcium phosphate ceramics driving osteogenesis through surface architecture, J. Biomed. Mater. Res. A (2014)http://dx.doi.org/10.1002/jbm.a.35272 (n/a–n/a). [14] A.M.C. Barradas, H. Yuan, C.A. van Blitterswijk, P. Habibovic, Osteoinductive biomaterials: current knowledge of properties, experimental models and biological mechanisms, Eur. Cell Mater. 21 (2011) 407–429 (discussion 429. http://www. ncbi.nlm.nih.gov/pubmed/21604242 (accessed June 20, 2014)). [15] P. Habibovic, T.M. Sees, M.A. van den Doel, C.A. van Blitterswijk, K. de Groot, Osteoinduction by biomaterials—physicochemical and structural influences, J. Biomed. Mater. Res. A 77 (2006) 747–762, http://dx.doi.org/10.1002/jbm.a.30712. [16] F. Matassi, A. Botti, L. Sirleo, C. Carulli, M. Innocenti, Porous metal for orthopedics implants, Clin. Cases Miner. Bone Metab. 10 (2013) 111–115 (http://www. pubmedcentral.nih.gov/articlerender.fcgi?artid=3796997&tool= pmcentrez&rendertype=abstract (accessed October 15, 2015)). [17] S. Bose, M. Roy, A. Bandyopadhyay, Recent advances in bone tissue engineering scaffolds, Trends Biotechnol. 30 (2012) 546–554, http://dx.doi.org/10.1016/j. tibtech.2012.07.005. [18] Tissue Engineering: Roles, Materials, and Applications, Nova Publishers, 2008 (http:// books.google.com/books?id=Aozd7XvWD7MC&pgis=1 (accessed July 7, 2014)). [19] C.M. Murphy, M.G. Haugh, F.J. O'Brien, The effect of mean pore size on cell attachment, proliferation and migration in collagen-glycosaminoglycan scaffolds for bone tissue engineering, Biomaterials 31 (2010) 461–466, http://dx.doi.org/10. 1016/j.biomaterials.2009.09.063. [20] R.A. Perez, H.-W. Kim, M.-P. Ginebra, Polymeric additives to enhance the functional properties of calcium phosphate cements, J. Tissue Eng. 3 (2012)http://dx.doi.org/ 10.1177/2041731412439555 (2041731412439555). [21] P.X. Ma, Scaffolds for tissue fabrication, Mater. Today. 7 (2004) 30–40, http://dx. doi.org/10.1016/S1369-7021(04)00233-0. [22] V.L. Tsang, S.N. Bhatia, Three-dimensional tissue fabrication, Adv. Drug Deliv. Rev. 56 (2004) 1635–1647, http://dx.doi.org/10.1016/j.addr.2004.05.001. [23] V. Karageorgiou, D. Kaplan, Porosity of 3D biomaterial scaffolds and osteogenesis, Biomaterials 26 (2005) 5474–5491, http://dx.doi.org/10.1016/j.biomaterials.2005. 02.002. [24] F. Krauss Juillerat, U.T. Gonzenbach, P. Elser, A.R. Studart, L.J. Gauckler, Microstructural control of self-setting particle-stabilized ceramic foams, J. Am. Ceram. Soc. 94 (2011) 77–83, http://dx.doi.org/10.1111/j.1551-2916.2010.04040.x. [25] J.R. Jones, L.L. Hench, Effect of surfactant concentration and composition on the structure and properties of sol–gel-derived bioactive glass foam scaffolds for tissue engineering, J. Mater. Sci. 38 (2003) 3783–3790, http://dx.doi.org/10.1023/A: 1025988301542. [26] Q. Lv, L. Nair, C.T. Laurencin, Fabrication, characterization, and in vitro evaluation of poly(lactic acid glycolic acid)/nano-hydroxyapatite composite microsphere-based scaffolds for bone tissue engineering in rotating bioreactors, J. Biomed. Mater. Res. A 91A (2009) 679–691, http://dx.doi.org/10.1002/jbm.a.32302. [27] Q.Z. Chen, I.D. Thompson, A.R. Boccaccini, 45S5 Bioglass®-derived glass–ceramic scaffolds for bone tissue engineering, Biomaterials 27 (2006) 2414–2425, http:// dx.doi.org/10.1016/j.biomaterials.2005.11.025. [28] R. Nazarov, H.-J. Jin, D.L. Kaplan, Porous 3-D scaffolds from regenerated silk fibroin, Biomacromolecules 5 (2004) 718–726, http://dx.doi.org/10.1021/bm034327e.

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939 [29] P. Gentile, M. Mattioli-Belmonte, V. Chiono, C. Ferretti, F. Baino, C. Tonda-Turo, et al., Bioactive glass/polymer composite scaffolds mimicking bone tissue, J. Biomed. Mater. Res. A 100A (2012) 2654–2667, http://dx.doi.org/10.1002/jbm.a. 34205. [30] S. del Valle, N. Miño, F. Muñoz, A. González, J.A. Planell, M.-P. Ginebra, In vivo evaluation of an injectable macroporous calcium phosphate cement, J. Mater. Sci. Mater. Med. 18 (2007) 353–361, http://dx.doi.org/10.1007/s10856-006-0700-y. [31] P. Chen, J. Tao, S. Zhu, Y. Cai, Q. Mao, D. Yu, et al., Radially oriented collagen scaffold with SDF-1 promotes osteochondral repair by facilitating cell homing, Biomaterials 39 (2015) 114–123, http://dx.doi.org/10.1016/j.biomaterials.2014.10.049. [32] W.J. Li, C.T. Laurencin, E.J. Caterson, R.S. Tuan, F.K. Ko, Electrospun nanofibrous structure: a novel scaffold for tissue engineering, J. Biomed. Mater. Res. 60 (2002) 613–621, http://dx.doi.org/10.1002/jbm.10167. [33] J.L. Simon, E.D. Rekow, V.P. Thompson, H. Beam, J.L. Ricci, J.R. Parsons, MicroCT analysis of hydroxyapatite bone repair scaffolds created via three-dimensional printing for evaluating the effects of scaffold architecture on bone ingrowth, J. Biomed. Mater. Res. A 85A (2008) 371–377, http://dx.doi.org/10.1002/jbm.a.31484. [34] A. Salerno, E. Di Maio, S. Iannace, P.A. Netti, Tailoring the pore structure of PCL scaffolds for tissue engineering prepared via gas foaming of multi-phase blends, J. Porous. Mater. 19 (2011) 181–188, http://dx.doi.org/10.1007/s10934-011-9458-9. [35] B.-H. Yoon, Y.-H. Koh, C.-S. Park, H.-E. Kim, Generation of large pore channels for bone tissue engineering using camphene-based freeze casting, J. Am. Ceram. Soc. 90 (2007) 1744–1752, http://dx.doi.org/10.1111/j.1551-2916.2007.01670.x. [36] S.-M. Lien, L.-Y. Ko, T.-J. Huang, Effect of pore size on ECM secretion and cell growth in gelatin scaffold for articular cartilage tissue engineering, Acta Biomater. 5 (2009) 670–679, http://dx.doi.org/10.1016/j.actbio.2008.09.020. [37] P.X. Ma, J.W. Choi, Biodegradable polymer scaffolds with well-defined interconnected spherical pore network, Tissue Eng. 7 (2001) 23–33, http://dx.doi.org/10. 1089/107632701300003269. [38] Q. Hou, D.W. Grijpma, J. Feijen, Preparation of interconnected highly porous polymeric structures by a replication and freeze-drying process, J. Biomed. Mater. Res. B Appl. Biomater. 67 (2003) 732–740, http://dx.doi.org/10.1002/jbm.b.10066. [39] J. Zhang, L. Wu, D. Jing, J. Ding, A comparative study of porous scaffolds with cubic and spherical macropores, Polymer 46 (2005) 4979–4985, http://dx.doi.org/10. 1016/j.polymer.2005.02.120. [40] L.J. White, V. Hutter, H. Tai, S.M. Howdle, K.M. Shakesheff, The effect of processing variables on morphological and mechanical properties of supercritical CO2 foamed scaffolds for tissue engineering, Acta Biomater. 8 (2012) 61–71, http://dx.doi.org/ 10.1016/j.actbio.2011.07.032. [41] F. Baino, C. Vitale-Brovarone, Mechanical properties and reliability of glass–ceramic foam scaffolds for bone repair, Mater. Lett. 118 (2014) 27–30, http://dx.doi.org/10. 1016/j.matlet.2013.12.037. [42] Q. Chen, F. Baino, S. Spriano, N.M. Pugno, C. Vitale-Brovarone, Modelling of the strength–porosity relationship in glass–ceramic foam scaffolds for bone repair, J. Eur. Ceram. Soc. 34 (2014) 2663–2673, http://dx.doi.org/10.1016/j.jeurceramsoc. 2013.11.041. [43] A. Dejaco, V.S. Komlev, J. Jaroszewicz, W. Swieszkowski, C. Hellmich, Micro CTbased multiscale elasticity of double-porous (pre-cracked) hydroxyapatite granules for regenerative medicine, J. Biomech. 45 (2012) 1068–1075, http://dx.doi. org/10.1016/j.jbiomech.2011.12.026. [44] D. Taylor, J.G. Hazenberg, T.C. Lee, The cellular transducer in damage-stimulated bone remodelling: a theoretical investigation using fracture mechanics, J. Theor. Biol. 225 (2003) 65–75, http://dx.doi.org/10.1016/S0022-5193(03)00222-4. [45] J.-S. Lee, H. Do Cha, J.-H. Shim, J.W. Jung, J.Y. Kim, D.-W. Cho, Effect of pore architecture and stacking direction on mechanical properties of solid freeform fabricationbased scaffold for bone tissue engineering, J. Biomed. Mater. Res. A 100 (2012) 1846–1853, http://dx.doi.org/10.1002/jbm.a.34149. [46] T. Serra, J.A. Planell, M. Navarro, High-resolution PLA-based composite scaffolds via 3-D printing technology, Acta Biomater. 9 (2013) 5521–5530, http://dx.doi.org/10. 1016/j.actbio.2012.10.041. [47] M. Yeo, C.G. Simon, G. Kim, Effects of offset values of solid freeform fabricated PCL–β-TCP scaffolds on mechanical properties and cellular activities in bone tissue regeneration, J. Mater. Chem. 22 (2012) 21636, http://dx.doi.org/10.1039/ c2jm31165h. [48] C. Wu, Y. Luo, G. Cuniberti, Y. Xiao, M. Gelinsky, Three-dimensional printing of hierarchical and tough mesoporous bioactive glass scaffolds with a controllable pore architecture, excellent mechanical strength and mineralization ability, Acta Biomater. 7 (2011) 2644–2650, http://dx.doi.org/10.1016/j.actbio.2011.03.009. [49] K.-W. Lee, S. Wang, L. Lu, E. Jabbari, B.L. Currier, M.J. Yaszemski, Fabrication and characterization of poly(propylene fumarate) scaffolds with controlled pore structures using 3-dimensional printing and injection molding, Tissue Eng. 12 (2006) 2801–2811, http://dx.doi.org/10.1089/ten.2006.12.2801. [50] J.M. Sobral, S.G. Caridade, R.A. Sousa, J.F. Mano, R.L. Reis, Three-dimensional plotted scaffolds with controlled pore size gradients: effect of scaffold geometry on mechanical performance and cell seeding efficiency, Acta Biomater. 7 (2011) 1009–1018, http://dx.doi.org/10.1016/j.actbio.2010.11.003. [51] J. Wu, Q. Zhao, J. Sun, Q. Zhou, Preparation of poly(ethylene glycol) aligned porous cryogels using a unidirectional freezing technique, Soft Matter 8 (2012) 3620, http://dx.doi.org/10.1039/c2sm07411g. [52] A. Arora, A. Kothari, D.S. Katti, Pore orientation mediated control of mechanical behavior of scaffolds and its application in cartilage-mimetic scaffold design, J. Mech. Behav. Biomed. Mater. 51 (2015) 169–183, http://dx.doi.org/10.1016/j.jmbbm. 2015.06.033. [53] S. Jia, L. Liu, W. Pan, G. Meng, C. Duan, L. Zhang, et al., Oriented cartilage extracellular matrix-derived scaffold for cartilage tissue engineering, J. Biosci. Bioeng. 113 (2012) 647–653, http://dx.doi.org/10.1016/j.jbiosc.2011.12.009.

937

[54] J.M. Cordell, M.L. Vogl, A.J. Wagoner Johnson, The influence of micropore size on the mechanical properties of bulk hydroxyapatite and hydroxyapatite scaffolds, J. Mech. Behav. Biomed. Mater. 2 (2009) 560–570, http://dx.doi.org/10.1016/j. jmbbm.2009.01.009. [55] S. Fiorilli, F. Baino, V. Cauda, M. Crepaldi, C. Vitale-Brovarone, D. Demarchi, et al., Electrophoretic deposition of mesoporous bioactive glass on glass–ceramic foam scaffolds for bone tissue engineering, J. Mater. Sci. Mater. Med. 26 (2015) 21, http://dx.doi.org/10.1007/s10856-014-5346-6. [56] C. Vitale-Brovarone, F. Baino, M. Miola, R. Mortera, B. Onida, E. Verné, Glass– ceramic scaffolds containing silica mesophases for bone grafting and drug delivery, J. Mater. Sci. Mater. Med. 20 (2009) 809–820, http://dx.doi.org/10.1007/s10856008-3635-7. [57] J.W. Lee, G. Ahn, J.Y. Kim, D.-W. Cho, Evaluating cell proliferation based on internal pore size and 3D scaffold architecture fabricated using solid freeform fabrication technology, J. Mater. Sci. Mater. Med. 21 (2010) 3195–3205, http://dx.doi.org/10. 1007/s10856-010-4173-7. [58] T.C. Lim, K.S. Chian, K.F. Leong, Cryogenic prototyping of chitosan scaffolds with controlled micro and macro architecture and their effect on in vivo neovascularization and cellular infiltration, J. Biomed. Mater. Res. A 94 (2010) 1303–1311, http://dx.doi.org/10.1002/jbm.a.32747. [59] F.J. O'Brien, B.A. Harley, I.V. Yannas, L.J. Gibson, The effect of pore size on cell adhesion in collagen-GAG scaffolds, Biomaterials 26 (2005) 433–441, http://dx.doi.org/ 10.1016/j.biomaterials.2004.02.052. [60] T. Mygind, M. Stiehler, A. Baatrup, H. Li, X. Zou, A. Flyvbjerg, et al., Mesenchymal stem cell ingrowth and differentiation on coralline hydroxyapatite scaffolds, Biomaterials 28 (2007) 1036–1047, http://dx.doi.org/10.1016/j.biomaterials.2006. 10.003. [61] G.-I. Im, J.-Y. Ko, J.H. Lee, Chondrogenesis of adipose stem cells in a porous polymer scaffold: influence of the pore size, Cell Transplant. 21 (2012) 2397–2405, http:// dx.doi.org/10.3727/096368912X638865. [62] S. Yamane, N. Iwasaki, Y. Kasahara, K. Harada, T. Majima, K. Monde, et al., Effect of pore size on in vitro cartilage formation using chitosan-based hyaluronic acid hybrid polymer fibers, J. Biomed. Mater. Res. A 81 (2007) 586–593, http://dx.doi. org/10.1002/jbm.a.31095. [63] D.J. Griffon, M.R. Sedighi, D.V. Schaeffer, J.A. Eurell, A.L. Johnson, Chitosan scaffolds: interconnective pore size and cartilage engineering, Acta Biomater. 2 (2006) 313–320, http://dx.doi.org/10.1016/j.actbio.2005.12.007. [64] C.G. Jeong, S.J. Hollister, Mechanical and biochemical assessments of threedimensional poly(1,8-octanediol-co-citrate) scaffold pore shape and permeability effects on in vitro chondrogenesis using primary chondrocytes, Tissue Eng. A 16 (2010) 3759–3768, http://dx.doi.org/10.1089/ten.TEA.2010.0103. [65] S.H. Oh, I.K. Park, J.M. Kim, J.H. Lee, In vitro and in vivo characteristics of PCL scaffolds with pore size gradient fabricated by a centrifugation method, Biomaterials 28 (2007) 1664–1671, http://dx.doi.org/10.1016/j.biomaterials.2006.11.024. [66] S.H. Oh, T.H. Kim, G. Il Im, J.H. Lee, Investigation of pore size effect on chondrogenic differentiation of adipose stem cells using a pore size gradient scaffold, Biomacromolecules 11 (2010) 1948–1955, http://dx.doi.org/10.1021/bm100199m. [67] S. Park, G. Kim, Y.C. Jeon, Y. Koh, W. Kim, 3D polycaprolactone scaffolds with controlled pore structure using a rapid prototyping system, J. Mater. Sci. Mater. Med. 20 (2009) 229–234, http://dx.doi.org/10.1007/s10856-008-3573-4. [68] J.X. Lu, B. Flautre, K. Anselme, P. Hardouin, A. Gallur, M. Descamps, et al., Role of interconnections in porous bioceramics on bone recolonization in vitro and in vivo, J. Mater. Sci. Mater. Med. 10 (1999) 111–120 (http://www.ncbi.nlm.nih.gov/ pubmed/15347932 (accessed June 20, 2014)). [69] F.P.W. Melchels, A.M.C. Barradas, C.A. van Blitterswijk, J. de Boer, J. Feijen, D.W. Grijpma, Effects of the architecture of tissue engineering scaffolds on cell seeding and culturing, Acta Biomater. 6 (2010) 4208–4217, http://dx.doi.org/10.1016/j. actbio.2010.06.012. [70] J.M. Kemppainen, S.J. Hollister, Differential effects of designed scaffold permeability on chondrogenesis by chondrocytes and bone marrow stromal cells, Biomaterials 31 (2010) 279–287, http://dx.doi.org/10.1016/j.biomaterials.2009.09.041. [71] A. Phadke, Y. Hwang, S.H. Kim, S.H. Kim, T. Yamaguchi, K. Masuda, et al., Effect of scaffold microarchitecture on osteogenic differentiation of human mesenchymal stem cells, Eur. Cell Mater. 25 (2013) 114–128 (discussion 128–9. http://www. ncbi.nlm.nih.gov/pubmed/23329467 (accessed June 20, 2014)). [72] Q. Fu, M.N. Rahaman, B.S. Bal, R.F. Brown, In vitro cellular response to hydroxyapatite scaffolds with oriented pore architectures, Mater. Sci. Eng. C 29 (2009) 2147–2153, http://dx.doi.org/10.1016/j.msec.2009.04.016. [73] S.R. Caliari, B.A.C. Harley, The effect of anisotropic collagen-GAG scaffolds and growth factor supplementation on tendon cell recruitment, alignment, and metabolic activity, Biomaterials 32 (2011) 5330–5340, http://dx.doi.org/10.1016/j. biomaterials.2011.04.021. [74] P. Yilgor, R.A. Sousa, R.L. Reis, N. Hasirci, V. Hasirci, 3D plotted PCL scaffolds for stem cell based bone tissue engineering, Macromol. Symp. 269 (2008) 92–99, http://dx.doi.org/10.1002/masy.200850911. [75] F. O'Brien, Influence of freezing rate on pore structure in freeze-dried collagen-GAG scaffolds, Biomaterials 25 (2004) 1077–1086, http://dx.doi.org/10.1016/S01429612(03)00630-6. [76] A. Salerno, D. Guarnieri, M. Iannone, S. Zeppetelli, P.A. Netti, Effect of micro- and macroporosity of bone tissue three-dimensional-poly(epsilon-caprolactone) scaffold on human mesenchymal stem cells invasion, proliferation, and differentiation in vitro, Tissue Eng. A 16 (2010) 2661–2673, http://dx.doi.org/10.1089/ten.tea. 2009.0494. [77] S.-W. Choi, Y. Zhang, Y. Xia, Three-dimensional scaffolds for tissue engineering: the importance of uniformity in pore size and structure, Langmuir 26 (2010) 19001–19006, http://dx.doi.org/10.1021/la104206h.

938

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939

[78] H.A. Declercq, T. Desmet, P. Dubruel, M.J. Cornelissen, The role of scaffold architecture and composition on the bone formation by adipose-derived stem cells, Tissue Eng. A 20 (2014) 434–444, http://dx.doi.org/10.1089/ten.TEA.2013.0179. [79] T.B.F. Woodfield, C.A. Van Blitterswijk, J. De Wijn, T.J. Sims, A.P. Hollander, J. Riesle, Polymer scaffolds fabricated with pore-size gradients as a model for studying the zonal organization within tissue-engineered cartilage constructs, Tissue Eng. 11 (2005) 1297–1311, http://dx.doi.org/10.1089/ten.2005.11.1297. [80] A. Penk, Y. Förster, H.A. Scheidt, A. Nimptsch, M.C. Hacker, M. Schulz-Siegmund, et al., The pore size of PLGA bone implants determines the de novo formation of bone tissue in tibial head defects in rats, Magn. Reson. Med. 70 (2013) 925–935, http://dx.doi.org/10.1002/mrm.24541. [81] F.M. Klenke, Y. Liu, H. Yuan, E.B. Hunziker, K.A. Siebenrock, W. Hofstetter, Impact of pore size on the vascularization and osseointegration of ceramic bone substitutes in vivo, J. Biomed. Mater. Res. A 85 (2008) 777–786, http://dx.doi.org/10.1002/ jbm.a.31559. [82] P. Duan, Z. Pan, L. Cao, Y. He, H. Wang, Z. Qu, et al., The effects of pore size in bilayered poly(lactide-co-glycolide) scaffolds on restoring osteochondral defects in rabbits, J. Biomed. Mater. Res. A (2013)http://dx.doi.org/10.1002/jbm.a.34683. [83] M. Belicchi, R. Cancedda, A. Cedola, F. Fiori, M. Gavina, A. Giuliani, et al., Some applications of nanotechnologies in stem cells research, Mater. Sci. Eng. B 165 (2009) 139–147, http://dx.doi.org/10.1016/j.mseb.2009.09.018. [84] J.R. Jones, G. Poologasundarampillai, R.C. Atwood, D. Bernard, P.D. Lee, Nondestructive quantitative 3D analysis for the optimisation of tissue scaffolds, Biomaterials 28 (2007) 1404–1413, http://dx.doi.org/10.1016/j.biomaterials.2006.11.014. [85] C. Renghini, A. Giuliani, S. Mazzoni, F. Brun, E. Larsson, F. Baino, et al., Microstructural characterization and in vitro bioactivity of porous glass–ceramic scaffolds for bone regeneration by synchrotron radiation X-ray microtomography, J. Eur. Ceram. Soc. 33 (2013) 1553–1565, http://dx.doi.org/10.1016/j.jeurceramsoc. 2012.10.016. [86] S. Kujala, J. Ryhänen, A. Danilov, J. Tuukkanen, Effect of porosity on the osteointegration and bone ingrowth of a weight-bearing nickel–titanium bone graft substitute, Biomaterials 24 (2003) 4691–4697 (http://www.ncbi.nlm.nih. gov/pubmed/12951012 (accessed June 20, 2014)). [87] O. Gauthier, J.M. Bouler, E. Aguado, P. Pilet, G. Daculsi, Macroporous biphasic calcium phosphate ceramics: influence of macropore diameter and macroporosity percentage on bone ingrowth, Biomaterials. 19 (1998) 133–139 (http://www. ncbi.nlm.nih.gov/pubmed/9678860 (accessed June 5, 2014)). [88] J.L. Simon, S. Michna, J.A. Lewis, E.D. Rekow, V.P. Thompson, J.E. Smay, et al., In vivo bone response to 3D periodic hydroxyapatite scaffolds assembled by direct ink writing, J. Biomed. Mater. Res. A 83 (2007) 747–758, http://dx.doi.org/10.1002/ jbm.a.31329. [89] A.I. Itälä, H.O. Ylänen, C. Ekholm, K.H. Karlsson, H.T. Aro, Pore diameter of more than 100 microm is not requisite for bone ingrowth in rabbits, J. Biomed. Mater. Res. 58 (2001) 679–683 (http://www.ncbi.nlm.nih.gov/pubmed/11745521 (accessed June 20, 2014)). [90] M.-C. von Doernberg, B. von Rechenberg, M. Bohner, S. Grünenfelder, G.H. van Lenthe, R. Müller, et al., In vivo behavior of calcium phosphate scaffolds with four different pore sizes, Biomaterials 27 (2006) 5186–5198, http://dx.doi.org/10. 1016/j.biomaterials.2006.05.051. [91] S.M.M. Roosa, J.M. Kemppainen, E.N. Moffitt, P.H. Krebsbach, S.J. Hollister, The pore size of polycaprolactone scaffolds has limited influence on bone regeneration in an in vivo model, J. Biomed. Mater. Res. A 92 (2010) 359–368, http://dx.doi.org/10. 1002/jbm.a.32381. [92] R.A. Perez, S.-J. Seo, J.-E. Won, E.-J. Lee, J.-H. Jang, J.C. Knowles, et al., Therapeutically relevant aspects in bone repair and regeneration, Mater. Today (2015)http://dx. doi.org/10.1016/j.mattod.2015.06.011. [93] R.A. Perez, J.-H. Kim, J.O. Buitrago, I.B. Wall, H.-W. Kim, Novel therapeutic core– shell hydrogel scaffolds with sequential delivery of cobalt and bone morphogenetic protein-2 for synergistic bone regeneration, Acta Biomater. 23 (2015) 295–308, http://dx.doi.org/10.1016/j.actbio.2015.06.002. [94] A. Hoppe, N.S. Güldal, A.R. Boccaccini, A review of the biological response to ionic dissolution products from bioactive glasses and glass–ceramics, Biomaterials 32 (2011) 2757–2774, http://dx.doi.org/10.1016/j.biomaterials.2011.01.004. [95] V. Mourino, J.P. Cattalini, Boccaccini, Metallic ions as therapeutic agents in tissue engineering scaffolds: an overview of their biological applications and strategies for new developments, J. R. Soc. Interface 9 (2012) 401–419, http://dx.doi.org/10. 1098/rsif.2011.0611. [96] B. Feng, Z. Jinkang, W. Zhen, L. Jianxi, C. Jiang, L. Jian, et al., The effect of pore size on tissue ingrowth and neovascularization in porous bioceramics of controlled architecture in vivo, Biomed. Mater. 6 (2011) 015007, http://dx.doi.org/10.1088/17486041/6/1/015007. [97] A.G. Mitsak, J.M. Kemppainen, M.T. Harris, S.J. Hollister, Effect of polycaprolactone scaffold permeability on bone regeneration in vivo, Tissue Eng. A 17 (2011) 1831–1839, http://dx.doi.org/10.1089/ten.TEA.2010.0560. [98] J.P. Li, P. Habibovic, M. van den Doel, C.E. Wilson, J.R. de Wijn, C.A. van Blitterswijk, et al., Bone ingrowth in porous titanium implants produced by 3D fiber deposition, Biomaterials 28 (2007) 2810–2820, http://dx.doi.org/10.1016/j.biomaterials.2007. 02.020. [99] P.J. Emans, E.J.P. Jansen, D. van Iersel, T.J.M. Welting, T.B.F. Woodfield, S.K. Bulstra, et al., Tissue-engineered constructs: the effect of scaffold architecture in osteochondral repair, J. Tissue Eng. Regen. Med. 7 (2013) 751–756, http://dx.doi. org/10.1002/term.1477. [100] K.M. Brouwer, W.F. Daamen, N. van Lochem, D. Reijnen, R.M.H. Wijnen, T.H. van Kuppevelt, Construction and in vivo evaluation of a dual layered collagenous scaffold with a radial pore structure for repair of the diaphragm, Acta Biomater. 9 (2013) 6844–6851, http://dx.doi.org/10.1016/j.actbio.2013.03.003.

[101] E.L.W. de Mulder, G. Hannink, N. Verdonschot, P. Buma, Effect of polyurethane scaffold architecture on ingrowth speed and collagen orientation in a subcutaneous rat pocket model, Biomed. Mater. 8 (2013) 025004, http://dx.doi.org/10.1088/17486041/8/2/025004. [102] T.M.G. Chu, D.G. Orton, S.J. Hollister, S.E. Feinberg, J.W. Halloran, Mechanical and in vivo performance of hydroxyapatite implants with controlled architectures, Biomaterials 23 (2002) 1283–1293 http://www.ncbi.nlm.nih.gov/pubmed/11808536 (accessed September 3, 2015). [103] E.L.W. de Mulder, G. Hannink, T.H. van Kuppevelt, W.F. Daamen, P. Buma, Similar hyaline-like cartilage repair of osteochondral defects in rabbits using isotropic and anisotropic collagen scaffolds, Tissue Eng. A 20 (2014) 635–645, http://dx. doi.org/10.1089/ten.TEA.2013.0083. [104] E.S. Sanzana, M. Navarro, M.-P. Ginebra, J.A. Planell, A.C. Ojeda, H.A. Montecinos, Role of porosity and pore architecture in the in vivo bone regeneration capacity of biodegradable glass scaffolds, J. Biomed. Mater. Res. A 102 (2014) 1767–1773, http://dx.doi.org/10.1002/jbm.a.34845. [105] X. Liu, M.N. Rahaman, Q. Fu, Bone regeneration in strong porous bioactive glass (13–93) scaffolds with an oriented microstructure implanted in rat calvarial defects, Acta Biomater. 9 (2013) 4889–4898, http://dx.doi.org/10.1016/j.actbio. 2012.08.029. [106] K.A. Hing, Bioceramic bone graft substitutes: Influence of porosity and chemistry, Int. J. Appl. Ceram. Technol. 2 (2005) 184–199, http://dx.doi.org/10.1111/j.1744– 7402.2005.02020.x. [107] D.D. Deligianni, N.D. Katsala, P.G. Koutsoukos, Y.F. Missirlis, Effect of surface roughness of hydroxyapatite on human bone marrow cell adhesion, proliferation, differentiation and detachment strength, Biomaterials 22 (2001) 87–96. [108] X.D. Zhu, H.J. Zhang, H.S. Fan, W. Li, X.D. Zhang, Effect of phase composition and microstructure of calcium phosphate ceramic particles on protein adsorption, Acta Biomater. 6 (2010) 1536–1541, http://dx.doi.org/10.1016/j.actbio.2009.10. 032. [109] A.L. Rosa, M.M. Beloti, R. van Noort, Osteoblastic differentiation of cultured rat bone marrow cells on hydroxyapatite with different surface topography, Dent. Mater. 19 (2003) 768–772, http://dx.doi.org/10.1016/S0109-5641(03)00024-1. [110] M. Espanol, R.A. Perez, E.B. Montufar, C. Marichal, A. Sacco, M.P. Ginebra, Intrinsic porosity of calcium phosphate cements and its significance for drug delivery and tissue engineering applications, Acta Biomater. 5 (2009) 2752–2762, http://dx. doi.org/10.1016/j.actbio.2009.03.011. [111] Y. Wang, E. Bella, C.S.D. Lee, C. Migliaresi, L. Pelcastre, Z. Schwartz, et al., The synergistic effects of 3-D porous silk fibroin matrix scaffold properties and hydrodynamic environment in cartilage tissue regeneration, Biomaterials 31 (2010) 4672–4681, http://dx.doi.org/10.1016/j.biomaterials.2010.02.006. [112] A. Mittal, P. Negi, K. Garkhal, S. Verma, N. Kumar, Integration of porosity and biofunctionalization to form a 3D scaffold: cell culture studies and in vitro degradation, Biomed. Mater. 5 (2010) 045001, http://dx.doi.org/10.1088/1748-6041/5/4/045001. [113] S. Yoshida, K. Hagiwara, T. Hasebe, a. Hotta, Surface modification of polymers by plasma treatments for the enhancement of biocompatibility and controlled drug release, Surf. Coat. Technol. 233 (2013) 99–107, http://dx.doi.org/10.1016/j. surfcoat.2013.02.042. [114] N.R. Washburn, K.M. Yamada, C.G. Simon, S.B. Kennedy, E.J. Amis, High-throughput investigation of osteoblast response to polymer crystallinity: influence of nanometer-scale roughness on proliferation, Biomaterials 25 (2004) 1215–1224, http://dx.doi.org/10.1016/j.biomaterials.2003.08.043. [115] N.E. Vrana, A. Dupret, C. Coraux, D. Vautier, C. Debry, P. Lavalle, Hybrid titanium/ biodegradable polymer implants with an hierarchical pore structure as a means to control selective cell movement, PLoS ONE 6 (2011), e20480http://dx.doi.org/ 10.1371/journal.pone.0020480. [116] U. Ripamonti, Bone induction in nonhuman primates. An experimental study on the baboon, Clin. Orthop. Relat. Res. 284–94 (1991) http://www.ncbi.nlm.nih. gov/pubmed/1864050 (accessed June 20, 2014). [117] C. Klein, K. de Groot, W. Chen, Y. Li, X. Zhang, Osseous substance formation induced in porous calcium phosphate ceramics in soft tissues, Biomaterials 15 (1994) 31–34 http://www.ncbi.nlm.nih.gov/pubmed/8161654 (accessed June 20, 2014). [118] H. Yamasaki, H. Sakai, Osteogenic response to porous hydroxyapatite ceramics under the skin of dogs, Biomaterials 13 (1992) 308–312 http://www.ncbi.nlm. nih.gov/pubmed/1318086 (accessed June 20, 2014). [119] M.P. Lutolf, J.a. Hubbell, Synthetic biomaterials as instructive extracellular microenvironments for morphogenesis in tissue engineering, Nat. Biotechnol. 23 (2005) 47–55, http://dx.doi.org/10.1038/nbt1055. [120] a. Howe, a.E. Aplin, S.K. Alahari, R.L. Juliano, Integrin signaling and cell growth control, Curr. Opin. Cell Biol. 10 (1998) 220–231. [121] F.G. Giancotti, Integrin signaling, Science 285 (80-) (1999) 1028–1033, http://dx. doi.org/10.1126/science.285.5430.1028. [122] G. Wei, P.X. Ma, Partially nanofibrous architecture of 3D tissue engineering scaffolds, Biomaterials 30 (2009) 6426–6434, http://dx.doi.org/10.1016/j.biomaterials.2009.08. 012. [123] J. Wang, H. Zhang, X. Zhu, H. Fan, Y. Fan, X. Zhang, Dynamic competitive adsorption of bone-related proteins on calcium phosphate ceramic particles with different phase composition and microstructure, J. Biomed. Mater. Res. B Appl. Biomater. 101 (2013) 1069–1077, http://dx.doi.org/10.1002/jbm.b.32917. [124] R.A. Perez, S. Del Valle, G. Altankov, M.-P. Ginebra, Porous hydroxyapatite and gelatin/hydroxyapatite microspheres obtained by calcium phosphate cement emulsion, J. Biomed. Mater. Res. B Appl. Biomater. 97 (2011) 156–166, http://dx.doi. org/10.1002/jbm.b.31798. [125] R.A. Perez, G. Altankov, E. Jorge-Herrero, M.P. Ginebra, Micro- and nanostructured hydroxyapatite-collagen microcarriers for bone tissue-engineering applications, J. Tissue Eng. Regen. Med. 7 (2013) 353–361, http://dx.doi.org/10.1002/term.530.

R.A. Perez, G. Mestres / Materials Science and Engineering C 61 (2016) 922–939 [126] R.A. Perez, T.-H. Kim, M. Kim, J.-H. Jang, M.-P. Ginebra, H.-W. Kim, Calcium phosphate cements loaded with basic fibroblast growth factor: delivery and in vitro cell response, J. Biomed. Mater. Res. A 101 (2013) 923–931, http://dx.doi.org/10. 1002/jbm.a.34390. [127] M. Espanol, I. Casals, S. Lamtahri, M.-T. Valderas, M.-P. Ginebra, Assessment of protein entrapment in hydroxyapatite scaffolds by size exclusion chromatography, Biointerphases 7 (2012) 37, http://dx.doi.org/10.1007/s13758-012-0037-7. [128] W. Cao, L.L. Hench, Bioactive materials, Ceram. Int. 22 (1996) 493–507, http://dx. doi.org/10.1016/0272-8842(95)00126-3. [129] J. De Groot, Carriers that concentrate native bone morphogenetic protein in vivo, Tissue Eng. 4 (1998) 337–341. [130] P. Habibovic, H. Yuan, C.M. van der Valk, G. Meijer, C.A. van Blitterswijk, K. de Groot, 3D microenvironment as essential element for osteoinduction by biomaterials, Biomaterials 26 (2005) 3565–3575, http://dx.doi.org/10.1016/j.biomaterials. 2004.09.056. [131] R.Z. LeGeros, Calcium phosphate-based osteoinductive materials, Chem. Rev. 108 (2008) 4742–4753, http://dx.doi.org/10.1021/cr800427g. [132] M.S. Lord, M. Foss, F. Besenbacher, Influence of nanoscale surface topography on protein adsorption and cellular response, Nano Today 5 (2010) 66–78, http://dx. doi.org/10.1016/j.nantod.2010.01.001. [133] C. Dai, H. Guo, J. Lu, J. Shi, J. Wei, C. Liu, Osteogenic evaluation of calcium/ magnesium-doped mesoporous silica scaffold with incorporation of rhBMP-2 by synchrotron radiation-based μCT, Biomaterials 32 (2011) 8506–8517, http://dx. doi.org/10.1016/j.biomaterials.2011.07.090. [134] C. Wu, W. Fan, J. Chang, Y. Xiao, Mesoporous bioactive glass scaffolds for efficient delivery of vascular endothelial growth factor, J. Biomater. Appl. 28 (2013) 367–374, http://dx.doi.org/10.1177/0885328212453635. [135] Y.F. Zhao, S.C.J. Loo, Y.Z. Chen, F.Y.C. Boey, J. Ma, In situ SAXRD study of sol–gel induced well-ordered mesoporous bioglasses for drug delivery, J. Biomed. Mater. Res. A 85 (2008) 1032–1042, http://dx.doi.org/10.1002/jbm.a.31545. [136] J.G. Dellinger, J.A.C. Eurell, R.D. Jamison, Bone response to 3D periodic hydroxyapatite scaffolds with and without tailored microporosity to deliver bone morphogenetic protein 2, J. Biomed. Mater. Res. Part A 76A (2006) 366–376, http://dx.doi. org/10.1002/jbm.a.30523. [137] J.R. Woodard, A.J. Hilldore, S.K. Lan, C.J. Park, A.W. Morgan, J.A.C. Eurell, et al., The mechanical properties and osteoconductivity of hydroxyapatite bone scaffolds with multi-scale porosity, Biomaterials 28 (2007) 45–54, http://dx.doi.org/10. 1016/j.biomaterials.2006.08.021. [138] B. Dorj, J.-E. Won, O. Purevdorj, K.D. Patel, J.-H. Kim, E.-J. Lee, et al., A novel therapeutic design of microporous-structured biopolymer scaffolds for drug loading and delivery, Acta Biomater. 10 (2014) 1238–1250, http://dx.doi.org/10.1016/j.actbio. 2013.11.002. [139] R.A. Perez, A. El-Fiqi, J.-H. Park, T.-H. Kim, J.-H. Kim, H.-W. Kim, Therapeutic bioactive microcarriers: co-delivery of growth factors and stem cells for bone tissue engineering, Acta Biomater. 10 (2014) 520–530, http://dx.doi.org/10.1016/j.actbio. 2013.09.042. [140] S. Radin, T. Chen, P. Ducheyne, The controlled release of drugs from emulsified, sol gel processed silica microspheres, Biomaterials 30 (2009) 850–858, http://dx.doi. org/10.1016/j.biomaterials.2008.09.066. [141] S. Kwon, R.K. Singh, R.A. Perez, E.A. Abou Neel, H.-W. Kim, W. Chrzanowski, Silicabased mesoporous nanoparticles for controlled drug delivery, J. Tissue Eng. 4 (2013)http://dx.doi.org/10.1177/2041731413503357 (2041731413503357). [142] R.A. Perez, M.-P. Ginebra, Injectable collagen/α-tricalcium phosphate cement: collagen–mineral phase interactions and cell response, J. Mater. Sci. Mater. Med. (2012)http://dx.doi.org/10.1007/s10856-012-4799-8. [143] R.A. Perez, H.-W. Kim, Core-shell designed scaffolds of alginate/alpha-tricalcium phosphate for the loading and delivery of biological proteins, J. Biomed. Mater. Res. A 101 (2013) 1103–1112, http://dx.doi.org/10.1002/jbm.a.34406. [144] D. Pastorino, C. Canal, M.-P. Ginebra, Drug delivery from injectable calcium phosphate foams by tailoring the macroporosity–drug interaction, Acta Biomater. 12 (2015) 250–259, http://dx.doi.org/10.1016/j.actbio.2014.10.031. [145] M.-P.P. Ginebra, C. Canal, M. Espanol, D. Pastorino, E.B. Montufar, Calcium phosphate cements as drug delivery materials, Adv. Drug Deliv. Rev. 64 (2012) 1090–1110, http://dx.doi.org/10.1016/j.addr.2012.01.008. [146] M.P. Ginebra, M. Espanol, E.B. Montufar, R.A. Perez, G. Mestres, New processing approaches in calcium phosphate cements and their applications in regenerative medicine, Acta Biomater. 6 (2010) 2863–2873, http://dx.doi.org/10.1016/j.actbio. 2010.01.036. [147] S.J. Lee, J.S. Choi, K.S. Park, G. Khang, Y.M. Lee, H.B. Lee, Response of MG63 osteoblast-like cells onto polycarbonate membrane surfaces with different

[148]

[149]

[150]

[151]

[152]

[153]

[154]

[155]

[156]

[157]

[158]

[159]

[160]

[161]

[162]

[163]

[164]

939

micropore sizes, Biomaterials 25 (2004) 4699–4707, http://dx.doi.org/10.1016/j. biomaterials.2003.11.034. Q. Zhang, Y. Jiang, Y. Zhang, Z. Ye, W. Tan, M. Lang, Effect of porosity on long-term degradation of poly (ε-caprolactone) scaffolds and their cellular response, Polym. Degrad. Stab. 98 (2013) 209–218, http://dx.doi.org/10.1016/j.polymdegradstab. 2012.10.008. Y. Takahashi, Y. Tabata, Effect of the fiber diameter and porosity of non-woven PET fabrics on the osteogenic differentiation of mesenchymal stem cells, J. Biomater. Sci. Polym. Ed. 15 (2004) 41–57 http://www.ncbi.nlm.nih.gov/pubmed/ 15027842 (accessed June 20, 2014). J. Rnjak-Kovacina, S.G. Wise, Z. Li, P.K.M. Maitz, C.J. Young, Y. Wang, et al., Tailoring the porosity and pore size of electrospun synthetic human elastin scaffolds for dermal tissue engineering, Biomaterials 32 (2011) 6729–6736, http://dx.doi.org/10. 1016/j.biomaterials.2011.05.065. F.A. Akin, H. Zreiqat, S. Jordan, M.B. Wijesundara, L. Hanley, Preparation and analysis of macroporous TiO2 films on Ti surfaces for bone-tissue implants, J. Biomed. Mater. Res. 57 (2001) 588–596 http://www.ncbi.nlm.nih.gov/pubmed/11553890 (accessed June 20, 2014). J. Isaac, J.-C. Hornez, D. Jian, M. Descamps, P. Hardouin, D. Magne, Beta-TCP microporosity decreases the viability and osteoblast differentiation of human bone marrow stromal cells, J. Biomed. Mater. Res. A 86 (2008) 386–393, http://dx.doi.org/10. 1002/jbm.a.31644. P. Kasten, I. Beyen, P. Niemeyer, R. Luginbühl, M. Bohner, W. Richter, Porosity and pore size of beta-tricalcium phosphate scaffold can influence protein production and osteogenic differentiation of human mesenchymal stem cells: an in vitro and in vivo study, Acta Biomater. 4 (2008) 1904–1915, http://dx.doi.org/10.1016/j. actbio.2008.05.017. Y. Wan, Y. Wang, Z. Liu, X. Qu, B. Han, J. Bei, et al., Adhesion and proliferation of OCT-1 osteoblast-like cells on micro- and nano-scale topography structured poly(L-lactide), Biomaterials 26 (2005) 4453–4459, http://dx.doi.org/10.1016/j. biomaterials.2004.11.016. K. Anselme, M. Bigerelle, B. Noel, E. Dufresne, D. Judas, a. Iost, et al., Qualitative and quantitative study of human osteoblast adhesion on materials with various surface roughnesses, J. Biomed. Mater. Res. 49 (2000) 155–166, http://dx.doi.org/10.1002/ (SICI)1097-4636(200002)49:2b155::AID-JBM2N3.0.CO;2-J. K. Anselme, P. Linez, M. Bigerelle, D. Le Maguer, A. Le Maguer, P. Hardouin, et al., The relative influence of the topography and chemistry of TiAl6V4 surfaces on osteoblastic cell behaviour, Biomaterials 21 (2000) 1567–1577. J. Lincks, B.D. Boyan, C.R. Blanchard, C.H. Lohmann, Y. Liu, D.L. Cochran, et al., Response of MG63 osteoblast-like cells to titanium and titanium alloy is dependent on surface roughness and composition, Biomater. Silver Jubil. Compend. 19 (2006) 147–160, http://dx.doi.org/10.1016/B978-008045154-1.50019-8. R.A. Gittens, T. McLachlan, R. Olivares-Navarrete, Y. Cai, S. Berner, R. Tannenbaum, et al., The effects of combined micron-/submicron-scale surface roughness and nanoscale features on cell proliferation and differentiation, Biomaterials 32 (2011) 3395–3403, http://dx.doi.org/10.1016/j.biomaterials.2011.01.029. G. Zhao, A.L. Raines, M. Wieland, Z. Schwartz, B.D. Boyan, Requirement for both micron- and submicron scale structure for synergistic responses of osteoblasts to substrate surface energy and topography, Biomaterials 28 (2007) 2821–2829, http://dx.doi.org/10.1016/j.biomaterials.2007.02.024. X. Zhu, J. Chen, L. Scheideler, R. Reichl, J. Geis-Gerstorfer, Effects of topography and composition of titanium surface oxides on osteoblast responses, Biomaterials 25 (2004) 4087–4103, http://dx.doi.org/10.1016/j.biomaterials.2003.11.011. R. Ikeda, H. Fujioka, I. Nagura, T. Kokubu, N. Toyokawa, A. Inui, et al., The effect of porosity and mechanical property of a synthetic polymer scaffold on repair of osteochondral defects, Int. Orthop. 33 (2009) 821–828, http://dx.doi.org/10. 1007/s00264-008-0532-0. K.A. Hing, B. Annaz, S. Saeed, P.A. Revell, T. Buckland, Microporosity enhances bioactivity of synthetic bone graft substitutes, J. Mater. Sci. Mater. Med. 16 (2005) 467–475, http://dx.doi.org/10.1007/s10856-005-6988-1. B.J. Story, W.R. Wagner, D.M. Gaisser, S.D. Cook, A.M. Rust-Dawicki, In vivo performance of a modified CSTi dental implant coating, Int. J. Oral Maxillofac. Implants 13 (1998) 749–757 (http://www.ncbi.nlm.nih.gov/pubmed/9857585 (accessed June 20, 2014)). C. Knabe, C. Koch, A. Rack, M. Stiller, Effect of beta-tricalcium phosphate particles with varying porosity on osteogenesis after sinus floor augmentation in humans, Biomaterials 29 (2008) 2249–2258, http://dx.doi.org/10.1016/j.biomaterials.2008. 01.026.