Patent Application
20220184276
The present disclosure relates to the field of medical technology, and in particular is directed to scaffolds for use in tissue engineering, a method for preparing scaffolds and uses of scaffolds for various medical applications in vivo and in vitro.
In tissue engineering, scaffolds have been proven their values for the repair of injury and regeneration of damaged tissues. A three-dimensional (3D) porous scaffold serves as a supporting framework for cell attachment, cell growth, and tissue regeneration. In order to facilitate soft tissue reconstruction into the constructs, scaffolds should closely mimic the biological environment surrounding the target tissue. To achieve this, scaffolds need to possess proper properties (mechanical, chemical and biological) so that they can deliver biological cues that cells can sense and respond to. Studies reveal that structure is a key aspect that influences the performance of a scaffold. Among the key structural parameters, porosity is one of the most important. For example, Perez et al [Perez et al, Materials Science and Engineering C, 61: 922-939, 2016] have reported the importance of both macro- and micro-porosity and the role of the overall architecture and morphology of the pores on the mechanical properties of printed scaffolds. The mechanical properties are also function of the design of the scaffolds [Slesarenko et al, Smart materials and structures, 26(3): 35053, 2017]. 3D devices with customized properties to the need of application can be manufactured by using 3D printing as method and by varying patterns, materials or combination of them [US2016374431 A1]. For example, the mechanical properties can be varied by printing grid-like pattern staggered or offset from one other [WO2014095872 A1].
The main problem when 3D-printing for medical applications today is that the scaffold becomes very stiff and the polymer degrades while printing if degradable polymers like poly(L-lactide), PLLA, poly(D-lactide), PDLA, poly(D,L-lactide), PDLLA, poly(meso-lactide), PmLA, poly(glycolide), PGA, poly(p-dioxanone) (PDX) and copolymers prepared by combinations of the relative monomers are printed. Poly(ε-caprolactone), PCL, does not degrade since the melting temperature is low and it has higher thermal stability compared to the other aliphatic polyesters, but the degradation time in vivo of high molecular weight PCL is too long to be adequate for medical applications.
Thus, there is a need in the art for improved scaffolds, which are not degraded during manufacture but which are nevertheless degradable when implanted in vivo in an individual, and which scaffolds have a decreased stiffness and an increased permeability to achieve improved cell attachment, cell growth, and tissue regeneration.
The above objective to provide improved scaffolds for medical applications is achieved by the present disclosure, which provides soft, degradable 3D-printable scaffolds by combining a gradient design and a staggered design of the scaffold.
More particularly, the present disclosure is directed to a scaffold for use in tissue engineering, comprising:
n layers, where n is an integer ≥4;
wherein each layer extends in an x-y-plane;
wherein each layer has a height (H) in a z-direction perpendicular to the x-y-plane;
wherein each layer comprises m volume-building components comprising scaffold material and being aligned in r rows in the x-y-plane, where m is an integer ≥2, where r is an integer ≥2;
wherein each volume-building component in any row of a layer is positioned at a distance (Dx) in the x-direction from any adjacent volume-building component in any adjacent row of the same layer;
wherein the m volume-building components of any group of layers are distributed at an angle (α) in the x-y-plane relative to the m volume-building components of any adjacent group of layers;
wherein the volume-building components are configured to allow biological cells to attach thereto;
characterised in that
the scaffold comprises a gradient distribution of volume-building components in the z-direction, wherein the distance (Dx) between any two adjacent groups of volume-building components of any group of layers is different from the distance (Dx) between any two adjacent groups of volume-building components of any adjacent group of layers;
and that
the scaffold comprises a staggered distribution of volume-building components of a layer in the x-y-plane, wherein each volume-building component in any row of a layer is positioned at a distance (Dysr) in the y direction from any adjacent volume-building component in the same row, and wherein each group of volume-building components in any group of rows of a layer is positioned at a distance (Dyar) in the y-direction from any adjacent group of volume-building components in any adjacent group of rows of the same layer;
wherein the height (H), the distance (Dx), the distance (Dysr), the distance (Dyar) and the angle (α) together define (or determine) a size and shape of pores in the scaffold, which pores are configured to allow biological cells to pass through the scaffold. Dysr is not always present. It presents when there are at least two volume-building components 5 in a row at a layer.
Further disclosed is a medical device for use in tissue engineering, comprising a scaffold as disclosed herein.
The present disclosure also provides a method for preparing a scaffold as described above by additive manufacturing, e.g. 3D-printing, in an apparatus configured for additive manufacturing.
The present disclosure is also directed to a method for in vivo tissue engineering, such as for breast reconstruction after mastectomy, preferably for supporting a breast prosthesis, or for reconstruction of tendon/ligament or muscle junction in an individual, comprising implanting a scaffold or a medical device presently disclosed into the individual.
Further provided herein is a use of the presently disclosed scaffold in an in vitro cell culture system, in an in vitro method for culturing of cells and/or in an in vitro method for regenerating tissue.
Also provided is a scaffold and a medical device for use in a method for in vivo tissue engineering.
Further disclosed is a novel degradable copolymer of ε-caprolactone and p-dioxanone, which can be printed without degradation and which is particularly suitable for use as scaffold material in the scaffold and method according to the present disclosure.
Preferred aspects of the present disclosure are further described below in the detailed description and in the dependent claims.
The present disclosure solves or at least mitigates the problems associated with existing scaffolds for use in medical applications by providing, as illustrated in
n layers 3, where n is an integer ≥4;
wherein each layer 3 extends in an x-y-plane;
wherein each layer 3 has a height (H) in a z-direction perpendicular to the x-y-plane;
wherein each layer 3 comprises m volume-building components 5 comprising scaffold material and being aligned in r rows 7 in the x-y-plane, where m is an integer ≥2, where r is an integer ≥2;
wherein each volume-building component 5 in any row 7 of a layer 3 is positioned at a distance (Dx) in the x-direction from any adjacent volume-building component 5 in any adjacent row 7 of the same layer 3;
wherein the m volume-building components 5 of any group of layers 3 are distributed at an angle (α) in the x-y-plane relative to the m volume-building components 5 of any adjacent group of layers 3; wherein the volume-building components 5 are configured to allow biological cells to attach thereto;
characterised in that
the scaffold 1 comprises a gradient distribution of volume-building components 5 in the z-direction, wherein the distance (Dx) between any two adjacent groups of volume-building components 5 of any group of layers 3 is different from the distance (Dx) between any two adjacent groups of volume-building components 5 of any adjacent group of layers 3;
and that
the scaffold 1 comprises a staggered distribution of volume-building components 5 of a layer 3 in the x-y-plane, wherein each volume-building component 5 in any row 7 of a layer 3 is positioned at a distance (Dysr) in the y direction from any adjacent volume-building component 5 in the same row 7, and wherein each group of volume-building components 5 in any group of rows 7 of a layer 3 is positioned at a distance (Dyar) in the y-direction from any adjacent group of volume-building components 5 in any adjacent group of rows 7 of the same layer 3;
wherein the height (H), the distance (Dx), the distance (Dysr), the distance (Dyar) and the angle (α) together define (or determine) a size and shape of pores in the scaffold 1, which pores are configured to allow biological cells to pass through the scaffold.
The shape and the size of pores can be precisely tailored by combining different values of Dx, Dysr, Dyar and α. Thereafter, the pores in the scaffold can be efficiently fabricated using additive manufacturing to possess desired characterization, such as shape and size distribution. This technology shows a huge advantage to other conventional fabrication methods (salt leaching, phase separation and gas foaming) in terms of high controllability and good reproducibility. The reason is that additive manufacturing uses data computer-aided-design (CAD) software or 3D object scanners to direct hardware to deposit material, layer upon layer, in precise geometric shape. For instance, the scaffold can be built from intercrossing struts deposited in the horizontal plane and form pores surrounded by struts with the aid of fused deposition modeling (FDM) technology. The complexity of the pore shape and size can be handled by FDM fabrication of a multi-layer design.
A significant advantage of the presently disclosed scaffold is that, as shown herein, the combination design of a gradient structure and a staggered structure can significantly improve the mechanical properties of a scaffold. Moreover, the reduction of the mechanical resistance of a scaffold to a given deformation can take place in both the z-direction (compression) and the x-y-plane (tension). Furthermore, the disclosed scaffold can possess a high porosity and a high permeability, which allows an efficient cell migration, as well as the transport of nutrients and waste.
It is to be understood that the phrase “layer extends in an x-y-plane” as used herein may be interpreted as “layer extends essentially in an x-y-plane”, meaning that any layer of a scaffold is able to extend in an x-y-plane but does not necessarily have to be completely flat. The scaffold material itself may have certain flexibility and softness, to make the scaffold adaptable to the forms of surrounding tissue when implanted in an individual. Additionally or alternatively, the scaffold may be printed on top of a soft textile or mesh and included in a medical implant which should likewise be formable in surrounding tissue in the body. The terms “x-y-plane” and “horizontal plane” may be used interchangeably herein.
Similarly, the phrase “layer has a height (H) in a z-direction perpendicular to the x-y-plane” as used herein may be interpreted as “layer has a height (H) in a z-direction essentially perpendicular to the x-y-plane”, meaning that the height of a layer is normally measured in a z-direction perpendicular to the x-y-plane, but the height of a layer does not necessarily have to be completely perpendicular to the x-y-plane. Again, this is due to the fact that the scaffold material itself may have certain flexibility and softness.
Herein, the term “volume-building component” includes any component which is capable of adding to the volume of the scaffold. A volume-building component may take any suitable form, and preferably has an elongated shape or form, such as a strand. Further, herein a volume-building component may comprise a plurality of volume-building elements joined together, provided that an element has a smaller size than a component. As a non-limiting example, a volume-building element may have the form of a droplet or sphere, and a plurality of such droplets or spheres may be joined together to form a strand. Some types of apparatuses for additive manufacturing are configured to deposit scaffold material in the form of elongated volume-building components, such as strands. In contrast, other types of apparatuses for additive manufacturing are configured to deposit scaffold material in the form of spheric volume-building elements, such as droplets or spheres, which may be adhered or fused together to form a chain of a predefined number of volume-building elements. The volume-building elements may for example be adhered or fused together by use of a suitable temperature, which is higher than the melting temperature of the scaffold material, while depositing the volume-building elements. Such a chain of volume-building elements may correspond to an elongated volume-building component in terms of its properties. The length of each chain of volume-building elements may be controlled. The chains of volume-building elements may be distributed in any desired pattern within one layer, such as a staggered distribution. Further, the distance Dx between any two adjacent chains of volume-building elements of a first group of layers may be controlled and may be different from the distance Dx between any two adjacent chains of volume-building elements of a second group of layers which is adjacent to the first group of layers.
In the context of the volume-building components being aligned in rows in the x-y-plane of any layer of the presently disclosed scaffold, the term “aligned” is intended to mean that essentially all volume-building components of a layer are arranged or positioned essentially in the same direction(s) or shape(s) in the x-y plane, but not necessarily in straight lines, meaning that the rows do not necessarily have to be straight lines. The volume-building components may for example be curved or otherwise shaped in the x-y-plane, resulting in rows having a curved or other shape. However, it is contemplated that essentially all volume-building components of a layer may be arranged in parallel to each other, meaning that all rows of a layer would be parallel to each other.
Throughout the present text, the term “adjacent” is defined as meaning next to, but not necessarily joined with, something else. Two components of the same type which are adjacent to each other do not have any component of the same type positioned inbetween them but may have a component of another type positioned inbetween them. For example, according to the present disclosure a first volume-building component and a second volume-building component which are adjacent (in a layer, or in a row) are positioned next to each other, without any third volume-building component positioned inbetween the first and second volume-building components. However, the first volume-building component and the second volume-building component normally have a certain distance between them in the x-direction and/or in the y-direction of the layer. According to another example of the present disclosure, a first layer and a second layer which are adjacent to each other are positioned next to each other in the scaffold, without any third layer positioned inbetween the first and second layers. Two adjacent layers may be joined (such as adhered, e.g. fused) or alternatively, there may be another type of component positioned inbetween them, in which case there is a certain distance between them in the z-direction. In the context of the present disclosure it is contemplated that for example a mesh or any type of supporting component may be positioned between any two adjacent layers of volume-building components. Thereby, the structure of the scaffold may be altered in many ways, which in turn results in different mechanical properties of the scaffold. In applications where a pliable scaffold is desirable, at least some of the layers of a scaffold preferably do not adhere to any adjacent layer, more preferably none of the layers adhere to any adjacent layer. In contrast, in applications where a scaffold with high rigidity is sought for, at least some of the layers of a scaffold, such as all layers, suitably adhere to adjacent layers, since adherence will increase the strength of the scaffold.
Herein, the term “group of volume-building components” is defined as being a group comprising at least one volume-building component, such as 1, 2, 3, 4, 5, 6, 7, 8, 9, 10 or more volume-building components. In an embodiment where the group comprises more than one volume-building component, the volume-building components of the group are adjacent to each other.
The term “group of layers” is to be interpreted as being a group comprising at least one layer, such as 1, 2, 3, 4, 5 or more layers. In an embodiment where the group comprises more than one layer, the layers of the group are adjacent to each other.
The term “group of rows” is to be interpreted as being a group comprising at least one row, such as 1, 2, 3, 4, 5 or more rows. In an embodiment where the group comprises more than one row, the rows of the group are adjacent to each other.
The m volume-building components of a any group of layers are distributed or oriented at an angle alpha (α) in the x-y-plane relative to the m volume-building components of any adjacent group of layers. In other words, the plurality of volume-building components of a first group of layers are aligned and are positioned or distributed or oriented at an angle α relative to the plurality of volume-building components of a second group of layers adjacent to the first group of layers. In this context, it is to be understood that a group of layers may suitably consist of one layer only, or alternatively may suitably consist of two layers. However, for some applications it is envisioned that it may be suitable to make a group consisting of more than two layers adjacent to each other have the same angle of orientation in relation to an adjacent group of layers. Further, the term “at an angle α in the x-y-plane” may alternatively be expressed as “at an angle α in a z-axis view”. Preferably said angle alpha is in a range of from >0° to 90°, such as 0.1°, 0.5°, 1°, 5°, 10°, 15°, 20°, 25°, 30°, 35°, 40°, 45°, 50°, 55°, 60°, 65°, 70°, 75°, 80°, 85°, or 90°. Changing the angle α results in a scaffold having a different stress response. For example, an angle α of 15° leads to a better stress distribution within the scaffold, and thus the scaffold has a lower stiffness compared to an angle α of 90°. It is to be understood that different degrees of the angle α are suitable for different applications in the field of medical technology.
The scaffold comprises a gradient distribution of volume-building components in the z-direction, wherein the distance Dx between any two adjacent groups of volume-building components of any group of layers is different from the distance Dx between any two adjacent groups of volume-building components of any adjacent group of layers. In other words, the distance Dx between (any two) adjacent groups of volume-building components of a first group of layers is different from the distance Dx between (any two) adjacent groups of volume-building components of a second group of layers which is adjacent to the first group of layers.
The gradient distribution of volume-building components in the z-direction may be achieved for example by any one or a combination of the following presently preferred configurations:
(a) The distance (Dx) between any two adjacent groups of volume-building components of a group of layers at a first end portion in the z-direction and/or at a second end portion in the z-direction of the scaffold is larger than the distance (Dx) between any two adjacent groups of volume-building components of a group of layers at an intermediate portion in the z-direction of the scaffold. This configuration results in scaffolds having a smaller pore size in the middle than at one or both ends in the z-direction of the scaffold. It is to be understood that this configuration may also include a repetitive pattern of alternating pore sizes in the z-direction of the scaffold.
(b) The distance (Dx) between any two adjacent groups of volume-building components of a group of layers at a first end portion in the z-direction of the scaffold is larger than the distance (Dx) between any two adjacent groups of volume-building components of a group of layers at a second end portion in the z-direction of the scaffold. This configuration results in scaffolds having a smaller pore size at one end and a larger pore size at the other end in the z-direction of the scaffold, or in other words results in scaffolds having a gradient of pore sizes from one end to the other end in the z-direction of the scaffold.
A third, currently less preferred, configuration is also contemplated, including a smaller pore size at both end portions than at an intermediate portion in the z-direction of the scaffold. It is to be understood that this configuration may also include a repetitive pattern of alternating pore sizes in the z-direction of the scaffold, meaning alternating decreased pore size and increased pore size, repeated periodically.
The minimum number of layers to achieve a gradient distribution of pore sizes in a scaffold is 4 layers, since a minimum of 2 layers are needed to form one layer of pores with a specific size, and thus 4 layers are required to form a gradient. Apart from this limitation in terms of a minimum number of layers to achieve a gradient distribution, it is to be understood that the properties of a scaffold are not dependent upon the number of layers in the scaffold. Suitably, the intended site or medical application of a scaffold may decide how many layers a specific scaffold should have.
Each volume-building component has a length L. The length L of each volume-building component affects the mechanical resistance to a deformation in the x-y-plane of a scaffold having a staggered distribution, and consequently also affects the properties of a medical implant comprising such a scaffold. For instance, the staggered structure results in a lower tensile modulus in the x-y-plane compared to a non-staggered structure. This length L may alternatively be called “continuous length” in the art. The length L depends on the intended site and medical application of a specific scaffold, and may vary within one scaffold. Volume-building components used in scaffolds described herein may suitably have a length L in a range of from about 1 mm to about 300 mm.
The height and width of a volume-building component may suitably be in a range of from about 0.05 mm to about 3 mm.
Further, the length or gap or distance between two adjacent volume-building components positioned in the same row or in adjacent rows of a layer also affects the mechanical resistance of a scaffold having a staggered distribution in a similar way as described above. This length or gap or distance may alternatively be denoted by the term “discontinuous length”.
Throughout the present text, the term “discontinuous length” is distinguished as a distance Dx in the x-direction when describing the distance between a volume-building component in any row of a layer and any adjacent volume-building component in any adjacent row of the same layer. Further, it is to be understood that in the phrase “the distance Dx between any two adjacent groups of volume-building components of any group of layers”, the distance Dx referred to is the distance Dx between a first volume-building component belonging to a first group of volume-building components, which first volume-building component is located at an edge of said first group, i.e. outermost in said first group, and which first volume-building component is located closest to a second volume-building component belonging to a second group of volume-building components, which second volume-building component is located outermost in said second group at an edge of said second group, which edge is located closest to said first group. The distance Dx depends on the intended site and medical application of a specific scaffold, and will vary within one scaffold to achieve a gradient distribution. However, normally the distance Dx may suitably be in a range of from about 0.2 μm to about 1500 μm.
Further, when describing the distance between a volume-building component in any row of a layer and any adjacent volume-building component in the same row, the term “discontinuous length” is distinguished as a distance Dysr (wherein “sr” stands for same row) in the y-direction. The distance Dysr depends on the intended site and medical application of a specific scaffold, and may vary within a scaffold. However, normally the distance Dysr may suitably be in a range of from about 0.2 μm to about 1500 μm.
Additionally, when describing the distance between a volume-building component in any row of a layer and any adjacent volume-building component in an adjacent row of the same layer, the term “discontinuous length” is distinguished as a distance Dyar (wherein “ar” stands for adjacent rows) in the y-direction. Further, it is to be understood that in the phrase “each group of volume-building components in any group of rows of a layer is positioned at a distance Dyar in the y-direction from any adjacent group of volume-building components in any adjacent group of rows of the same layer”, the distance Dyar referred to is the distance Dyar between a first volume-building component belonging to a first group of rows of a layer, which first volume-building component is located at an edge of said first group, i.e. outermost in said first group, and which first volume-building component is located closest to a second volume-building component belonging to a second group of rows of the same layer, which second volume-building component is located outermost in said second group at an edge of said second group, which edge is located closest to said first group. The distance Dyar depends on the intended site and medical application of a specific scaffold, and may vary within a scaffold. However, normally the distance Dyar may suitably be in a range of from about 0.2 μm to about 1500 μm.
The scaffold comprises a staggered distribution of volume-building components in the x-y-plane, wherein each volume-building component in any row of a layer is positioned at a distance (Dysr) in the y-direction from any adjacent volume-building component in the same row, and wherein each group of volume-building components in any group of rows of a layer is positioned at a distance (Dyar) in the y-direction from any adjacent group of volume-building components in any adjacent group of rows of the same layer. In other words, the horizontal position of a group of volume-building components of a layer is offset relative to the horizontal position of any adjacent group of volume-building components of said layer.
It should be clarified that a staggered distribution of volume-building components in a scaffold can be achieved through different configurations in terms of ‘discontinuous length’. For example, a scaffold comprises only Dysr, or Dyar, or a combination of Dysr and Dyar to form a staggered distribution of volume-building components. In this regard, the minimum number of volume-building components at one layer to achieve a staggered distribution in a scaffold is varying. Specifically, a minimum of 2 volume-building components are needed to form an offset (Dyar) from one row to another at one layer and a minimum of 4 volume-building components are needed to form an offset (Dysr) from one row to another, wherein there are at least 2 volume-building components in one row.
More particularly, the staggered distribution of volume-building components, wherein the horizontal position of a group of volume-building components of a layer is offset relative to the horizontal position of any adjacent group of volume-building components of said layer, may be achieved for example if the length L of each of a group of volume-building components of a layer is smaller than the width W of said layer, i.e. L<Wx or L<Wy. Alternatively, or additionally, the staggered distribution may be achieved, even if L >Wx or L >Wy, if each of a group of volume-building components of a layer is distributed, positioned or formed, e.g. curved or diagonally positioned, within the width Wx and/or Wy of the layer, such that each of the group of volume-building components extends a distance in the x-direction and/or a distance in the y-direction, which distance is shorter than the width Wx and Wy, respectively, of said layer.
Each layer of a scaffold has a height H in the z-direction, i.e. along a z-axis. The height H of each layer may be the same for all layers of a scaffold or may be different for different layers of a scaffold. The height of a layer depends on the intended site and medical application of the scaffold, and may vary within a scaffold. However, normally the height H may suitably be in a range of from 0.05 to 0.4 mm.
The height of a complete scaffold, comprising a plurality of layers, depends on the intended site and medical application of the scaffold. However, normally the height of a complete scaffold may suitably be in a range of from about 0.2 mm to about 10 mm. In a currently preferred medical application of the scaffold, namely reconstruction of a breast, the height of a scaffold may preferably be from about 0.2 mm to about 3 mm, wherein the maximum height corresponds to the thickness of the so-called Cooper's ligament in the breast.
Each layer of a scaffold may be defined as having a width Wx in the x-direction and a width Wy in the y-direction. The width Wx of a layer may be equal to the width Wy of the layer, in which case the layer has a circular shape. Alternatively, the width Wx of a layer may be different from the width Wy of the layer, resulting in any type of shape of the layer. Further, it is to be understood that different layers of a scaffold may have different widths and different shapes, depending on the intended application in each specific case, for example the type and/or shape of a defect to be treated, or the type of tissue to be treated. The widths Wx and Wy of a layer may suitably be in a range of from about 1 mm to about 300 mm.
The scaffold according to the present disclosure may comprise scaffold material which is resorbable or degradable in vivo.
The scaffold material used to prepare the scaffold according to the present disclosure may comprise a polymer, a copolymer or a combination of different polymers and/or copolymers. Preferably, the polymer(s) and/or copolymer(s) comprise one or more monomers selected from a group consisting of glycolide, lactide (D- or L-, meso or racemic mixture), trimethylene carbonate, ε-caprolactone, p-dioxanone (i.e. paradioxanone), β-butyrolactone and 1,5-dioxepan-2-one, or combinations thereof. More preferably, the scaffold material comprises a copolymer of ε-caprolactone and p-dioxanone as disclosed herein. Such a copolymer preferably has a random distribution of ε-caprolactone and p-dioxanone units along the polymeric chain and a higher content of ε-caprolactone than p-dioxanone. The copolymer is prepared by bulk ring-opening polymerization of a mixture of the two monomers and has preferably a melting point between 60° C. and 25° C. The copolymer forms films with a Young's Modulus between 50 and 400 MPa and elongation to break between to 200 and 2000%, and hydrolytic degradation time between 8-24 months depending on the composition, but much faster than PCL.
The microstructure of a polymer or copolymer, e.g. random or blocky microstructure, influences the physical properties of the scaffold. Thus, the microstructure is one parameter which may be varied to achieve scaffolds which are suitable for different medical applications. It is to be understood that for the purpose of preparing a scaffold by additive manufacturing according to the presently disclosed method, the microstructure of the polymer may be of any type, such as random, or blocky, as long as the parameters of the method for preparing the scaffold are tailored to the type of microstructure desired for each medical application.
The present disclosure further provides a medical device for use in tissue engineering, comprising a scaffold as described in detail above.
The medical device may further comprise any suitable type of support material, such as a non-woven or mesh material, for example in the form of a sheet of mesh, or a homogeneous film.
Alternatively or in addition, the medical device as described above may comprise a plurality of scaffolds as described above, wherein the plurality may comprise 2, 3, 4, 5, 10, 15, 20 or more scaffolds. Such scaffolds included in the same medical implant may have identical or different properties. As a non-limiting example, the porosity of two scaffolds within the same medical implant may be identical or different. Other properties which may vary between scaffolds within a medical implant are the pore size, pore structure, interconnectivity (i.e. open porosity), and mechanical properties.
It is to be understood that the presently disclosed scaffold, as well as the presently disclosed medical device, are preferably configured to be implanted in an individual. Non-limiting examples include that the scaffold or medical device may be configured to be sutured or stapled or otherwise fastened to surrounding tissue, and/or may be pliable to conform to the surrounding tissue.
The present disclosure further provides a method for preparing a scaffold, as illustrated schematically in the flow chart of
(A) providing a scaffold material;
(B) preparing a scaffold from the scaffold material in an apparatus configured for additive manufacturing comprising a material source configured to deposit volume-building components of scaffold material, comprising:
It is to be understood that by performing the above-described method, each volume-building component in any row of a layer of the scaffold will be positioned at a distance (Dx) in the x-direction from any volume-building component in any adjacent row of the same layer. Further, the m volume-building components of any group of layers will be distributed at an angle (α) in the x-y-plane relative to the m volume-building components of any adjacent group of layers. Also, the scaffold will comprise a gradient distribution of volume-building components in the z-direction, wherein the distance (Dx) between any two adjacent groups of volume-building components of any group of layers is different from the distance (Dx) between any two adjacent groups of volume-building components of any adjacent group of layers, and the scaffold will comprise a staggered distribution of volume-building components of a layer in the x-y-plane, wherein each volume-building component in any row of a layer is positioned at a distance (Dysr) in the y direction from any adjacent volume-building component in the same row, and wherein each group of volume-building components in any group of rows of a layer is positioned at a distance (Dyar) in the y-direction from any adjacent group of volume-building components in any adjacent group of rows of the same layer.
The term “apparatus configured for additive manufacturing” is intended to mean an apparatus configured to build layers of volume-building components of scaffold material. Additive manufacturing uses data computer-aided-design (CAD) software or 3D object scanners to direct hardware to deposit material, layer upon layer, in precise geometric shapes. Such an apparatus may alternatively be called for example a rapid prototyper, a manufacturing processer, or a 3D printer. The apparatus comprises a base plate and a material source. The term “material source” may alternatively be called head of a material feeding system, head of material extrusion system, nozzle, or printer head. It is to be understood that different types of apparatuses configured for additive manufacturing may be used to perform the presently disclosed method for preparing a scaffold. Non-limiting examples of suitable apparatus for use in the presently disclosed method are apparatus configured for filament extrusion and fused deposition modeling (FDM), fused filament fabrication (FFF), 3D dispensing, 3D bioplotting extrusion and material jetting (such as freeformer), or Directed Energy Deposition (DED).
Further, the person skilled in the art understands that parameters and conditions suitable to apply in the method for preparing a scaffold, e.g. time, temperature, dimensions of volume-building components, distances Dx, Dysr, Dyar and height H, will vary depending on the intended site or medical application for a specific scaffold. Dimensions of volume-building components, distances Dx, Dysr, Dyar and height H have been described in detail above. Further below is given two non-limiting examples of a method for preparing a scaffold comprising a presently preferred scaffold material.
The distance Dyar between two adjacent volume-building components, positioned in two adjacent rows of a layer, may be as long as the entire width Wx or Wy of the layer, meaning that after the material source has deposited the last volume-building component of a row, the material source moves back the entire width Wx or Wy of the layer before starting to deposit the first volume-building component of the subsequent row.
It is common knowledge in the art that in many apparatuses configured for additive manufacturing, either the material source, the base plate or both of them may be moved. This means that instead of moving the material source a certain distance Dx, Dysr, and/or Dyar before depositing a subsequent volume-building component or moving the material source a distance (Dh) in the z-direction, equal to the height (H) of a layer, before forming a nsubsequent layer, the base plate may be moved an equal distance in the opposite direction to achieve the same structure of the scaffold.
As mentioned above in relation to the presently disclosed scaffold, adjacent layers may be made to adhere to each other but not necessarily. Conditions sufficient to adhere a subsequent layer to the preceding layer will vary depending on for example the type of scaffold material used.
The scaffold material used in the presently disclosed method may comprise a polymer, a copolymer or a combination of different polymers and/or copolymers, preferably wherein the polymer(s) and/or copolymer(s) comprise one or more monomers selected from a group consisting of glycolide, lactide (D- or L-, meso or racemic mixture), trimethylene carbonate, ε-caprolactone, p-dioxanone, β-butyrolactone and 1,5-dioxepan-2-one, or combinations thereof, more preferably wherein the scaffold material comprises a copolymer of ε-caprolactone and p-dioxanone. The scaffold material used preferably has a melting point below or equal to 60° C. and a thermal degradation starting 200° C. above the melting point. Such material preferably is a random copolymer of ε-caprolactone and p-dioxanone and has higher content of ε-caprolactone than p-dioxanone.
When using the presently preferred scaffold material, i.e. comprising a copolymer of ε-caprolactone and p-dioxanone, when performing the method for preparing a scaffold, the method may comprise the following non-limiting conditions. Fused deposition modeling (FDM) at temperature between 140-220° C. and speed 10-20 mm s−1 of extruded filament of the copolymer at 125° C. and 9 rpm can be used for scaffold preparation, or 3D bioplotting of polymer pellets at temperature in the range 140-160° C. and 8.5 bar of pressure.
According to yet another aspect, the present disclosure provides a method for in vivo tissue engineering comprising implanting the presently disclosed scaffold or medical device into an individual. In particular, such a method may be applied for soft tissue engineering, such as adipose tissue engineering, and more particularly for breast reconstruction, such as after mastectomy, wherein the scaffold may preferably be used for supporting a breast prosthesis. Without wishing to be bound by theory, it is believed that the scaffold may be configured to mimic the so-called Cooper's ligament which is a supporting structure, like a frame, in the mammalian breast. Other specific applications of such a method for in vivo tissue engineering include reconstruction of tendon, ligament or muscle junction in an individual.
Further provided is the presently disclosed scaffold for use in a method for in vivo tissue engineering as described in detail above. Also provided is the presently disclosed medical device for use in a method for in vivo tissue engineering as described in detail above.
The present disclosure is also directed to in vitro uses of the presently disclosed scaffold, such as, but not limited to, a use of the presently disclosed scaffold in an in vitro cell culture system, in an in vitro method for culturing of cells and/or in an in vitro method for regenerating tissue.
Herein, devices and scaffolds “comprising” one or more recited elements may also include other elements not specifically recited. The term “comprising” includes as a subset “consisting essentially of” which means that the device or scaffold has the components listed without other features or components being present. Likewise, methods “comprising” one or more recited steps may also include other steps not specifically recited. Further, the singular “a” and “an” shall be construed as including also the plural.
For a scaffold with a given porosity, the mechanical properties can be tailored by altering the strand alignments. Moreover, internal structure has a great impact on crucial factors for soft tissue regeneration, such as transport of nutrient, adhesion of cells and deposition of matrix. A computer-aided design method with the introduction of computational simulation technologies, namely finite element analysis (FEA) and computational fluid dynamics (CFD), is suitable to gain a priori knowledge of a scaffold architecture in terms of mechanical properties and transport behaviors. A further step is to fabricate designed 3D scaffolds and validate the computational analysis against experimental results. In this study, FDM technology was utilized to fabricate scaffolds with present designs and thereafter validate numerical analysis. During printing, filaments are melted through the heated extruder and deposited on the printer bed layer by layer to create an object. This method is the most cost-effective way of producing thermoplastic objects and prototypes. The objective of this study was to investigate the effect of designed structures on the mechanical properties and transport behaviors of 3D porous scaffolds by means of numerical and experimental procedures.
2.1 Scaffold Structure Design
In this study, ten 3D structural designs are presented in Table 1. For all designs, a 9-layer scaffold was constructed by repeating strands with a constant thickness of 0.4 mm. The difference in scaffold designs were controlled by tailoring strand orientation (SO), strand space (SS) and strand thickness (ST). The original length of each strand was 12 mm. After assigning SO, SS and ST values to a given design, a boolean operation was carried out to obtain a scaffold with a cylindrical shape. The radius was 5 mm. To be specific, three types of structures were examined, namely normal (N), gradient (G) and gradient+stagger (GS). A normal structure is referred to as SO changes from one layer to another while keeping a constant SS at the same layer. Five different SO values of 10, 15, 30, 45 and 90 degree combines with a constant SS value of 0.4 mm (0.8 and 1.2 mm for 90 degree as well) were considered for a normal structure. Moreover, a varying SS was presented in a gradient structure, which possesses a large SS at the first end and the second end portion and a small SS at the middle portion in the perpendicular direction of a scaffold. Two SO values of 15 and 90 degree are combined with a varying SS value of 1.2-0.8-0.4-0.8-1.2 from the first end to the second end of a scaffold. Furthermore, a combination of a staggered structure and a gradient structure was formulated. In the staggered structure, a shorter length of 8 mm was assigned to each strand. Meanwhile, strands were aligned in a staggered manner by offsetting their horizontal positions to adjacent strands. As a consequence, strands at each layer were separated from the boundary of the scaffold with a certain distance. Only one SO value of 15 degree combines with a varying SS values of 1.2-0.8-0.4-0.8-1.2 mm for gradient+stagger structure. The theoretical porosity of each scaffold is listed in Table 1, which is estimated by following equation,
p=(1 −Vsolid)/Vtotal (1)
where p is the porosity, Vsolid is the solid volume of a scaffold structure and Vtotal is total volume of the occupied by a scaffold structure.
2.2 Computational Analysis
The 3D model was built up in ABAQUS 6.14 and used for FEA and CFD simulations (
E=F
R/(Aε) (2)
where E is the effective compressive/tensile modulus, FR is the reaction force, A is the cross sectional area, and ε is the strain.
In CFD model, the scaffold was inserted into a virtual bioreactor that includes one inlet and one outlet. The velocity and the pressure throughout the model were obtained by solving the continuity and Navier-Stokes equations using ABAQUS CFD solver. Water was considered as the working fluid, which has a density of 1000 kg·m−3 and a dynamic viscosity of 0.001 kg·m−1·s−1. The diameter of the inlet and the outlet was 0.002 m. A velocity of 10−4 m·s−1 was applied for the inlet, which guarantees Reynolds number Re <1 in the range of validity of Darcy's law. A zero pressure was defined as the boundary condition at the outlet. Moreover, no-slip wall conditions were assigned for the walls of the scaffold and the virtual bioreactor. The average pressure values at the top and the bottom of the scaffold were extracted and used to determine of the permeability coefficient from the Darcy's law,
k=QμL/(AΔP) (3)
where k is the permeability, Q is the fluid flow rate, L is the length of the scaffold, A is the cross sectional area, and ΔP is the pressure drop from the top to the bottom of the scaffold.
Thereafter, all the permeability coefficient obtained were normalized by the reference permeability coefficient, which is predicted from a separate CFD analysis of a 100% porosity model.
2.3 Fabrication of PCL Scaffolds
Scaffolds were printed using the fifth generation MarkerBot Replicator Desktop 3D Printer (Stratasys, USA). The 3D printer uses filaments of 1.75 mm in diameter and it was fitted with a nozzle of 0.40 mm. All CAD files were made in Abaqus then sliced and converted to .makerbot file using MakerBot Print. The PCL filaments (wt. %) used were commercially available (3D4Makers, Netherlands) and used as received. The slicing distance was set to 0.4 mm and the printing speed to 10 mm−1. No infill, roof, floors, rafts or additional support structures were used. The retraction distance was set to zero and one shell was used, of each scaffold designs ten structures were produced each time each time at a printing temperature of 100° C. with no additional cooling of the printing plate, the excess polymer was cut away from the structures.
2.4 Mechanical Testing
Compressive testing of printed scaffolds (N15, N90, G90, G15 and GS15) were conducted to determine mechanical properties using the Instron 5566 tensile testing instrument at room temperature and 50% humidity using a 10 kN load cell after have been equilibrated for 24 hours. The scaffolds were compressed at a deformation speed of 0.35 mm−1, the equivalent of approximately 10% of the height of the scaffolds. The force and the displacement were recorded throughout the testing, and thereafter converted to a stress-strain curve. Each scaffold was compressed at least 1.5 mm and the compressive modulus reported for scaffolds was defined as the slope of the initial linear region in the stress-strain curve (n=7).
2.5 Microcomputed Tomography (μ-CT) Characterization of Scaffolds
All scaffolds were scanned using Skyscan 1172 system (Bruker Micro-CT, Kontich, Belgium) at 9 μm spatial resolution. For all scans, the operation voltage setting was 40 kV, the current utilized was 250 μA, and no filter was employed. A step size of 0.4° and frame averaging of 2. For reconstructions, NRecon (Bruker Micro CT, Kontich, Belgium) were carried out using full cone beam Feldkamp reconstruction algorithm with greyscale limits defined automatically (Need to insert REF-check Mohammeds paper). The microstructure of 3D printed scaffolds of PCL were analyzed from the CT images using CTAn (CT analyser, ver. 1.18, SkyScan). For each scaffold, region of interest (ROI) was selected and thresholding of these ROI was carried out using automatic Otsu thresholding. Surface area, total porosity, interconnectivity, mean pore size and pore size distribution were identified (n=3). 3D images were generated using CTVox (CTVox version 2.7, SkyScan) software
3.1 Mechanical Properties
The compressive and tensile moduli of five designed structure obtained from FEA simulations are shown in
A comparison of experimental and computational compressive moduli for scaffolds (N90, N15, G90, G15, GS15) are shown in
3.2 Transport Behaviors
The study of mass transport behavior inside the scaffold was carried out in a virtual bioreactor model. The permeability was extracted from CFD simulations as shown in
3.3 Characterization of 3d Printed Scaffolds
As shown in
In the present study, three types of structural designs (normal, gradient, gradient+stagger) were presented for soft tissue engineering with the assistance of computational simulation methods and FDM printing technology. Among these designs, a gradient combined with a stagger structure shows the best performance in terms of mechanical properties and mass transport. With this design, the scaffold possesses the lowest compressive/tensile modulus (approximately 0.23% of pure PCL) and the highest permeability. The computational compressive moduli for scaffolds were validated and in good agreement with measurements from compression testing. The largest difference was 16.5%. In addition, porosity effect is not the only key factor to influence the properties of a scaffold. As evidenced by results in this study, structure effect does play a role particularly in aspects of mechanical properties and mixing behaviors. Moreover, μCT analysis of printed scaffolds shows that FDM printing technology is capable of producing high-quality scaffolds from the designed structure. Lastly, the used method and findings in this study provides a priori knowledge of how architecture influence the performance of scaffolds.
Poly(ε-caprolactone), PCL, is a hydrophobic, semicrystalline resorbable polymer widely used for scaffolds manufacturing for tissue engineering applications. PCL is a prime candidate for melt-extrusion 3D printing of scaffolds. The low melting point, around 60° C., is an advantage when the polymer is melt-processed, indeed, low processing temperature can be used avoiding chain degradation, which typically occurs when other degradable polymers are melt-processed. Moreover, PCL has higher thermal stability and a wider processing window but the high crystallinity, the hydrophobicity and the long degradation time (up to years) are serious limitations of PCL-based scaffolds. Poly(p-dioxanone) (PDX) is instead a semicrystalline poly(ether ester) that exhibits flexibility and pliability, as well as greater hydrophilicity and a faster degradation rate than PCL (6 months), but really poor thermal stability.
The aim of this study was to develop by bulk copolymerization of CL with p-dioxanone (DX) a copolymer suitable for the fabrication of scaffolds that could be processed in a manner similar to that for PCL, but have more rapid degradation rates than PCL while maintaining a certain degree of crystallinity, a low Tm and high thermal stability.
2.1. Bulk Ring-Opening Polymerization
Copolymerization reaction of CL and DX were performed in bulk, using stannous octoate, Sn(Oct)2, as catalyst and ethylene glycol as initiator. The ratio catalyst to initiator was 1:17 and the ratio monomers to initiator was calculated in order to have a Mn,th=80000 g mol−1. The reaction temperature was 130° C., while the copolymerization time was in the range 24-48 h.
2.2. Film Preparation
Films of copolymers containing increasing amounts of DX were prepared with a thickness of ˜200 μm by casting a polymer solution in CHCl3 at a concentration of 100 g L−1 in a 9 cm diameter Petri dish. Smaller size films were then cut into the desired shape for mechanical, contact angle and degradation tests.
2.3. Degradation Tests
Degradation tests were performed over 120 days during which data point samples were collected at 4 times and each data point was analyzed in triplicate. Small films with a dimeter of 11 mm (30-40 mg weight) were dried in vacuum to a constant weight before tests were performed. The samples were immersed in vials containing 30 mL of sodium phosphate buffer, pH 7.4 and were maintained at 37° C. The buffer solution was replaced every 3 weeks. Films were withdrawn from the incubation medium at scheduled times, washed carefully with distilled water, dried to a constant weight to calculate the mass loss (%) and then analyzed by SEC and NMR.
2.4. Copolymer and Film Characterization
NMR spectra of polymer and film samples were obtained in CDCl3 at room temperature using a Bruker Avance 400 spectrometer (1H: 400.13; 13C: 100.62 MHz). The resonances and coupling constants are reported in ppm (δ) and Hz (J), respectively. 1H NMR spectra were referenced to the residual solvent proton at δ 7.26 ppm; 13C NMR spectra were referenced to the 13C signal of CDCl3 at δ 77.16 ppm. Spectra were recorded using Bruker TopSpin v2.1 software. Data processing was performed using MestReNova v9.0.0 software. Molecular weights (Mn and Mw) and dispersity (Ð) were measured by size exclusion chromatography (SEC). The measurements were performed at 35° C. on a GPCMAX system equipped with an RI detector and three columns: one guard column (PLgel 5 μm Guard) and two linear columns (PLgel 5 μm Mixed-D), using CHCl3 as the eluent (0.5 mL min-1). Narrow molecular weight polystyrene standards were used as standards and flow rate fluctuations were corrected using toluene as an internal standard.
Inherent viscosity (IV) measurements were performed on a Ubbelohde viscometer (capillary number 0, Capillary constant 0.009904) with the temperature set at 25° C. For the analysis, 10 mg of polymer sample were dissolved in hexafluoroisopropanol (HFIP) and filtered using a 0.45 mm PTFE membrane. The results are expressed as the mean value of triplicate measurements.
Melting points (Tm) and enthalpies of fusion (ΔHm) of the polymeric samples were measured by differential scanning calorimetry (DSC) using aluminum pans and a Mettler Toledo DSC 1 Instruments apparatus that was calibrated with indium. Measurements were performed under nitrogen flow at a heating rate of 10° C. min-1 from −30 to 120° C. DSC data are reported for the second heating cycle. The degree of crystallinity, Xc was calculated from the DSC results by considering an enthalpy of fusion for an infinitely large PCL crystals of 136.1 J g−1.
Thermogravimetric analysis (TGA) was performed in an O2 atmosphere at a flowrate of 80 mL min−1 in a temperature range from 25° C. to 500° C. at a heating rate of 10° C. min−1 using a TGA/DSC 1 Instrument apparatus (Mettler Toledo, USA). The results are reported as the temperature at which 5% mass loss occurred (T5%).
The tensile mechanical properties of films (7 cm length, 0.45 cm width, 200 μm thickness) were measured at room temperature using an Instron 5944 tensile testing machine equipped with pneumatic grips. The free clamped length was 36 mm. A 500 N load cell was used and the cross-head speed was 500% mm min−1 based on the initial length. All tensile properties are reported as the average of five measurements. Prior to analysis the specimens were conditioned for 24 h at 25° C. and 50% humidity.
Contact angle measurements were performed on films having a thickness of ˜200 μm that were prepared by casting from a CHCl3 solution at a polymer concentration of 100 g L−1 by placing a droplet of distilled water (MilliQ) on the surface and the angle was then estimated using CAM 2008 software. The reported values are the average of three measurements. Prior to analysis, the specimens were conditioned for 24 h at 25° C. and 50% humidity.
Different polymerization runs were conducted by increasing the amount of DX in the feed from 0 to 20 mol %. Sn(Oct)2 was used as the catalyst and ethylene glycol as the initiator and the reactions were performed in bulk at 130° C. (Table 2).
aMolar percentage of monomer in the feed.
bRatio is calculated to produce a Mn, th ~80 kg mol − 1.
cDetermined from the 1H NMR spectra (CDCl3, 400 MHz) of the crude copolymerization mixture.
dDetermined from the 1H NMR spectra (CDCl3, 400 MHz) of the purified copolymer.
eDetermined by SEC (CHCl3, versus polystyrene standards).
fDetermined by viscosimetry in HFIP at 25° C.
The composition is close to the feed in all the cases and high conversion was achieved for both monomers. The experimental number average molecular weight values, Mn, determined by SEC, were close to the theoretical values. The inherent viscosity of the polymer samples was also determined and it tends to decrease with increasing DX content. The copolymer microstructure consisted of PCL chains containing randomly distributed, but isolated, DX units, by increasing the DX content, the length of the segments comprising consecutive CL units decreased.
All of the synthesised copolymers were semicrystalline materials. The Tm decreased from 60.8 to 29.3° C. by increasing the DX content from 0 to 20 mol %, which is a consequence of the increased disorder along the polymer chains. A perfect linearly correlation was observed between the Tm and the amount of CL in the copolymer. The degree of crystallinity varied in the range 35-19%, depending on the composition (Table 3).
aDetermined from the 1H NMR spectra (CDCl3, 400 MHz) of the purified copolymer.
bDSC data reported from the cooling run.
cDSC data reported from the second heating run.
dCalculated from DSC data assuming a ΔH°PCL = 136.1 J g−1.
The thermal stability of the synthesized copolymers was evaluated by TGA analysis. The temperature difference between the start of the degradation and the melting point was approximately 235-245° C. for all of the polymers, indicating that, similar to PCL, the copolymers can be melt processed in a wide range of temperatures. The copolymers retained a certain degree of crystallinity and showed good thermal stability comparable to that of PCL. Copolymer films showed stress-strain curves that are typical of soft and flexible thermoplastic materials. The mechanical properties are summarized in Table 4.
Contact angle measurements performed on solvent-cast films showed that by increasing the amount of DX in the copolymer, the hydrophobicity was reduced. PCL and the copolymer containing 8 mol % DX have contact angles greater than 90°, therefore films made from these polymers have a hydrophobic surface; increasing the amount of DX to 13 and 20 mol %, causes the contact angles to decrease to 86° and 80°, respectively, thus proving that DX moieties reduce the hydrophobicity of the polymer surface.
Hydrolytic degradation study performed in vitro showed that The Mn values of the copolymer films decrease linearly with the degradation time; they are reduced by 50-80%, depending on the composition, after 120 days. Copolymer films containing 13 and 20 mol % DX have a Mn that is lower than 20 kg mol-1 after 120 days. No variations in the Ð values were observed. The copolymer can reach a Mn of 3000 g mol−1 in 8-9 months. 3000 g mol−1 is the limiting molecular weight, below which PCL oligomers are soluble and can diffuse out of the matrix, undergo intracellular resorption and lead to erosion of the implant.
Degradable polymers with low melting point suitable for the manufacturing of medical devises, mainly scaffolds for tissue engineering, that degrade in the human body, or other biological environments, and that can be melt processed at low temperature, thus, avoiding chain degradation during melt processing were developed. The polymers are random copolymers of ε-caprolactone (CL) and p-dioxanone (DX) having a content of DX from 5 to 20 mol %. Such materials combine the good processability and thermal stability of poly(caprolactone), PCL, with the hydrophilicity, good degradability and faster degradation of poly(p-dioxanone), PDX. Indeed, despite the low melting point, below 60° C., the copolymers have high degradation temperature and thermal stability comparable to that of (PCL), and increased hydrophilicity as well as faster degradation time than PCL.
It is to be understood that the present disclosure is not restricted to the above-described exemplifying embodiments thereof and that several conceivable modifications of the present disclosure are possible within the scope of the following claims.
| Number | Date | Country | Kind |
|---|---|---|---|
| 1950395-2 | Mar 2019 | SE | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/SE2020/050176 | 2/17/2020 | WO | 00 |
