Team:KCL UK/Scaffold Design

Our Scaffold Design:

In order to be able to start realizing our scaffold we had to define some requirements based on our knowledge of the human spinal cord. These can be subdivided into three categories, such as mechanical properties, biodegradability and biocompatibility. Following such parameters would then let us choose on the appropriate material, shape and architecture for the scaffold.

After an injury, the spinal cord is subject to the application of a variety of loads and stresses, leading to a significant (temporary) change in its geometry. Mechanical deformation can be the cause of functional and anatomical deficiency of spinal axons. For this reason, it is crucial that the mechanical properties of the proposed scaffold will match the ones of the spinal cord, in order for it not to deform under stress nor to be altered because of compression. Although, another way of achieving desired mechanical strength would be by crosslinking and optimising the material composition. Some key biomechanical properties, related to the spinal cord, to be considered when implementing the realisation of the scaffold are the following:

  • The mechanical properties of the extracellular matrix in tissue, for example stiffness, can highly alter cell behaviour, hence it is important to avoid a mechanical mismatch between engineered and host tissue, as this might be the cause of additional damage.
  • Spinal cord tissue is soft and shows viscoelastic properties, under deformation it exhibits a combination of both viscous (fluid) and elastic (solid) behaviours; for these reasons, ideally, the scaffold would require mechanical flexibility.
  • Additionally, stress-relaxation behaviour of the spinal cord is difficult to quantify, as it appears that CNS tissue stiffness would be likely to increase with age.

Biodegradability is another key requirement to consider while designing our scaffold. This is not only related to the choice of the material to use but also to the regeneration of axons, which would be the main achievement of our implementation. Biodegradable polymers are in fact beneficial both because they provide structural support for axon regeneration and because they can also act as a conduit for time-dependent delivery of therapeutic agents (if required). Ideally, the degradation rate of the material should follow the timescale of axonal regeneration, so that the mechanical integrity of the scaffold can be maintained throughout the regenerative phase. This parameter though can be regulated through external factors such as tuning the stiffness of the scaffold, the degree of acetylation and by altering the length of hydrolytically degradable units within the polymer crosslink. Furthermore, the scaffold must be bioresorbable, being safely reabsorbed in the body and not releasing any unnatural product when degrading, as this might become harmful or toxic to the body. The implemented design has to be also biocompatible, as it should elicit no immune response.

Following the outline of the primary requirements for our model, we would then be able to use this knowledge to explore different material and choose the most appropriate. The main difference in material was between natural and synthetic polymers. Both of these biomaterials have pros and cons to be considered:

  • Synthetic materials have many benefits, such as a better mechanical strength, higher machinability and degradation rates which result to be more controllable compared to natural materials. However, such materials are hostile for cell adhesion and exhibit rather lower biological activity and higher hydrophobicity.
  • Naturally obtained materials, show instead better biological achievements in cell proliferation and differentiation

The variable that outweighed the others when selecting a biomaterial was mechanical strength, as explained earlier this is in fact one of the key requirements of the scaffold. Furthermore, when deciding on how to shape the biomaterial the main achievement would be to create a design from which axonal regeneration would benefit. Analysing the architecture, two different aspects, macro-architecture and micro-architecture. Macro-architecture focuses on the external shape of the scaffold, which is considered to be more important of its surface area. When considering different shapes various consideration need to be made:

  • Cylinder, tube and channel implants has shown not to encourage tissue bridging
  • Open-path designs on the other hand, seems to enhance spinal cord regeneration as well as allowing for extension of myelinated fibers both externally and internally, along the length of the defect.

Micro-architecture instead, relies on pores. Porous scaffolds play an important role in the field of tissue engineering, as they control cell functions and guide the formation of new tissue. Moreover, porosity improves the connectivity between neurons, enhancing axons and glial cell growth. As it has been shown in many studies, a higher porosity improves axon growth, but if an optimal percentage is not used, this might create additional problems to the structure and functionality of the scaffold. Finally, when considering the general architecture of the implementation, its orientation needs to be considered as well. This should mimic the physiological arrangement of axons which are positioned longitudinally.

To summarise, the requirements for the realisation of an ideal scaffold are:

  • Possessing mechanical properties which matches the ones of the spinal cord
  • Being biodegradable, bioresorbable and biocompatible
  • Having a shape that will assure the accomplishment of its objectives
  • Optimal pores size

We validated our scaffold requirements as part of our integrated human practises and from these validated scaffold requirements, we created a scaffold specification sheet. We ensured each engineering decision taken was in line with our scaffold specification sheet and hence our human practises.

Upon assessing that a scaffold was the best approach to treat spinal cord injury patients, we began researching to find the optimal biomaterial to be used for our scaffold. There are two categories of potential biomaterials for scaffolds including synthetic biomaterials and natural biomaterials. Through our research into both synthetic and natural scaffolds, we identified cell adhesion as an area where scaffolds were not yet optimal for treating spinal cord injury, despite their promise and advantages over the other novel approaches. Our team split into sub-teams, with a one subteam designing a bio-adhesive protein coating to use synthetic biology to design an approach facilitating axonal regrowth along the scaffold, and a scaffold design subteam focusing on designing the personalised scaffold approach providing structural support and to fill the site of the glial scar in SCI patients. The scaffold design team initially conducted literature-based research looking into non-toxic, biocompatible natural and synthetic biomaterial candidates for our scaffold. The scaffold requirements specification sheet was central to choosing material. As the specification sheet was designed with human practises in mind, it was essential that the chosen material was in line with the scaffold requirements.

Natural Materials

Naturally derived materials have been identified as suitable candidates and utilised in scaffolds in pre-clinical research due to the relatively high biological activity, lower hydrophobicity and as they are more favourable for cellular adhesive, when compared with synthetic materials. We initially conducted literature research into naturally derived polymer candidates prior to researching synthetic material alternatives, keeping our scaffold requirements in mind to ensure human practises has been considered before making engineering design decisions (Zhang et al., 2019).


Collagen meets the scaffold requirements by being porous, biocompatible, with a low immunogenicity. Collagen is broken due by a catabolic process involving collagenase, which degrades the collagen fibres, making collagen a biodegradable and resorbable candidate. Playing an important role in maintaining the structural integrity of the extra cellular matrix, collagen provides structural support of issues which would be beneficial for axonal regrowth. Collagen has been used in murine models to bridge a transected spinal cord and promoted modest neurite growth, supporting its suitability to a SCI scaffold. Despite the modest neurite growth, high concentrations of collagen can impede neurite growth, which is detrimental to axonal regrowth. Collagen has poor mechanical properties, which is not in line with our scaffold material requirements. In order to overcome this, collagen scaffolds can be combined with synthetic polymers in a blend to optimise the mechanical properties(Dong and Lv, 2016).


Alginate has been identified as a suitable material to create biocompatible scaffolds to facilitate axonal regrowth. After initial promising pre-clinical results showing alginate enhanced neural sprouting and reduced inhibitory cues after an injury, future in vitro studied reported the inhibition of growth of dorsal root ganglion neurones and alters the phenotype of a variety of CNS cells, reducing their metabolic activities. The atyptical cellular appearance and activity surrounding the scaffold suggested that alginate was an unsuitable candidate for our scaffold. Alginate has poor mechanical properties for a structural supportive scaffolds(Zhang et al., 2018).


Chitosan is biocompatible, biodegradable and non-toxic. Chitosan has been identified as a promising material for scaffold design and can have varying mechanical properties, molecular weights and porosity, resulting in the ability to optimise mechanical properties. Research has been conducted into the rate of chitosan degradation, facilitating convenient modelling of the chitosan scaffolds to providing evidence that the scaffolds will not degrade too rapidly. Chitosan has been identified as a good candidate material to blend with other materials because of the cellular affinity chitosan has. Chitosan can also be combined with synergistic therapies including drug delivery systems, which could benefit the SCI treatment approach. Despite this, chitosan is influenced by the degree of acetylation which can influence the ability of cells, included neurons, to adhere and migrate along the surface. Our scaffold requires axons to adhere and regrow through the scaffold pores and along the scaffold. Through our research, we concluded that chitosan alone rather than in a blend is not an optimal material candidate as the mechanical properties are poor in a wet state, such as within the spinal cord, becoming brittle so chitosan was only considered as a promising candidate for a blended material scaffold(Madihally and Matthew, 1999).


Agarose, a linear polysaccharide, has been validated in vivo and has been found to elicit no severe adverse reactions. Agarose matches the scaffold requirements with regards to the mechanical properties, which can be engineered to match the mechanical properties of the spinal cord. Agarose has been implemented in vivo in the spinal cord and has shown to be stable for sufficient time after implementation. Despite the benefits, agarose alone is unable to promote robust growth of neural tissues when compared to collagen and fibrin(Zarrintaj et al., 2018).

Fibrin and Fibronectin

Alternative natural material candidates were also explored including fibrin and fibronectin, yet these materials were optimal for hydrogel scaffolds and hence did not offer the structural support with the optimal porous structure which is beneficial for facilitating axonal regrowth in chronic-stage cervical spinal cord injury patients(Ahmed et al., 2008).

Synthetic Biomaterials

Through initial literature research before consulting experts, we explored the suitability of promising synthetic material candidates, by evaluating if the materials would facilitate the scaffold meeting the scaffold requirements. When compared with natural materials, synthetic materials have better mechanical strength, higher machinability and a more optimal degradation rate. Synthetic materials have been employed as the only material or as part of a blend in pre-clinical research for a variety of therapeutic approaches and are biocompatible, eliciting a minimal immune response (Zhang et al., 2019).

Polysaccharide-based polymers including PCL

PCL is an FDA-approved non-toxic, bioresorbable polymer. PCL has been identified as an excellent candidate choice for scaffolds, supported by the suitable mechanical properties and the slow degradation rate of the biomaterial. Immunological studies have shown that PCL elicits a minimum immune response and chronic inflammatory response. The pore orientation and porosity can be controlled with PCL, facilitating the ability to optimise the scaffold’s microarchitecture. PCL is a relatively inexpensive material choice, which would assist in creating a more accessible treatment option than a more expensive material candidate(Siddiqui et al., 2018).

Poly(ethylene glycol) (PEG)

PEG is a suitable biocompatible scaffold candidate and is well tolerated by the body, having been successfully used in a variety of applications including cell cultures, cosmetics and drug delivery systems. PEG has been suggested to have some neuroprotective effects and has been shown to reunite and fuse transected cells processes, yet has been found to resist cell adhesion and protein absorption(Sakiyama-Elbert et al., 2012).

Aliphatic polyesters (PLA, PGA)

Aliphatic polyesters including PLA and PGA are biodegradable and biocompatible, with good mechanical properties. PGLA can be formed from PGLA with the addition of lactic acid repeats which can module the mechanical properties and physical properties. PLGA has shown to support axonal regrowth in vitro. Aliphatic polymers have a fast degradation rate, which is not optimal for a scaffold which needs to remain in the spinal cord for long enough to facilitate sufficient axonal regrowth with structural support. PGLA has a more optimal degradation rate when compared to PLA and PGA, but in vitro the degradation rate has been shown to be as little as 2-10 weeks, which can result in an acid microenvironment upon degradation which could be a contributing factor to preventing axonal regeneration(Du et al., 2014).

Poly(2-hydroxyethyl methyacrylate) (pHEMA)

PHEMA consists of a crosslinked network of hydrophilic co-polymers, that can provide a medium for cell attachment and growth in vivo. As a non-degradable, biocompatible and non-toxic material which has been successfully used for therapeutic approaches, pHEMA was considered. Despite the established benefits, pHEMA was discounted as a suitable scaffold candidate as pHEMA is non-biodegradable and non-resorbable, with sub-optimal micro-architecture for a porous scaffold(Sakiyama-Elbert et al., 2012).

Graphene and Carbon-Nanotubes (CNT)

Graphene and CNT were explored as potential novel materials for scaffold use, but the expensive cost of the material would mean the treatment we were designing would be expensive and hence inaccessible(Ahn et al., 2015).


Through literature research, we investigated the feasibility of a composite blend material for our scaffold, combining a natural and a synthetic polymer. Composite materials are chosen when existing biomaterials are not optimal to meet the clinical requirements. We investigated a PCL and chitosan blend and a PCL and collagen blend to enhance the mechanical properties of natural material choices. Through keeping our human practises in mind and ensuring we were choosing the most optimal material candidate with patient safety in mind, we coupled the literature research with discussions of the suitability of the most promising biomaterial candidates with experts.

We interviewed Dr. Riehle, the Director of the Centre of Cell Engineering at the University of Glasgow, and discussed our material choices. Dr. Riehle informed us that collagen is a relatively expensive biomaterial choice. We decided to discount this material for use in a blend or alone in our scaffold as our aim is to produce an accessible treatment for SCI, which the cost would prevent. We also considered a chitosan blend, with the successful use of chitosan in drug delivery systems being an advantage of the natural material, which we considered using alongside PCL in a composite scaffold. We were informed by Dr. Riehle that chitosan is likely to prove immunogenic in vivo and hence, was not an optimal scaffold material. This led us to instead focus on PCL. Similarly to chitosan, PCL can also act as a drug delivery system, allowing for future treatment developments of the PCL scaffold treatment.

We interviewed Dr. Koffler, an Assistant Professor of Neurosciences at the University of California San Diego who has undertaken research into the utilisation of scaffolds for the treatment of SCI. Through our discussion with Dr. Koffler about the suitability for using PCL in the scaffold, we finalised our material choice as PCL and solidified that this was the optimal material candidate to fit our scaffold requirements.

Research: Background Reading & Defining the Problem

Having decided upon a scaffold, arguably the two most vital features of which to consider would be the micro and macro architectures. The macro-architecture is concerned with the shape of the scaffold. A variance in shape can lead to varying degrees of success when it comes to axonal regeneration, defect length and astrocyte migration. Most scaffolds that have been developed have tubes or channels implemented. However, these two designs do not take into account the white and grey matter; grey matter is softer than white matter (Budday et al., 2015). Furthermore, research into how macro-architecture affects regeneration is scarce. Wong et al. (2008) proposed five scaffold designs: tube, channel, cylinder, open path with core and open path without core. Our sub-team will perform a literature review of the paper by Wong et al. (2008) to evaluate the proposed scaffold designs. In addition to this, we would simulate the scaffolds to compare their behaviour when a gravitational load is applied. This is an important aspect to consider because we want to see how each scaffold would perform in vivo, thus allowing us to select the most optimal design. However, we are not able to implant these in vivo, so our in silico results will only be an approximation.

Imagine: Brainstorm & Evaluate

From Wong et al. (2008) the best designs were the open path designs with the open path with core slightly outperforming the open path without core. Axonal regeneration was observed in both of these implants and the core in the open path with core design allowed for the white matter tract to be supported.

Table 1. Evaluation of the 5 scaffold designs proposed by Wong et al. (2008)

Advantages Disadvantages
Cylinder Defect length doubled, during the same time period.
Large scar and cyst formation.
No astrocyte migration.
Tube Defect length doubled, during the same time period.
Large scar and cyst formation.
Increased wall thickness of the prosthesis (more polymer mass) caused more swelling during polymer degradation, and in combination with smaller diameter guides caused more nerve compression, thus, resulting in overall decreased peripheral nerve regeneration.
Robust axonal regrowth with multiple channels of 450-660 microns when seeded with schwann cells. Defect length doubled, during the same time period.
Large scar and cyst formation.
Rim fibrous tissue surrounding nerve of regenerated neurons (larger channel diameter= more tissue and less axon regenerated.)
Open-path with core Supports white matter tracts.
Allows extension of myelinated fibers along the length of the defect exterior and inside the scaffold.
Maintained defect size over 3 months.
Axonal regeneration observed.
Open-path without core Supports white matter tracts.
Allows extension of myelinated fibers along the length of the defect exterior and inside the scaffold.
Maintained defect size over 3 months.
Axonal regeneration observed.
Only showed bundles of fibers on the outside.

Having obtained our conclusion from the literature review, we wanted to see what results in silico research would give. Observing the stresses, strains, displacement and applied force of each scaffold provides a useful insight in choosing the correct design. We proposed to create each design as a CAD file and simulate them in Autodesk Inventor with the NASTRAN extension.

Design & Build: Develop and Prototype Solution

The following prototypes were designed in Autodesk Inventor.
Figure 1. Macro-architecture designs implemented in Autodesk Inventor, based on the designs by Wong et al. (2008). (a) Open path without core, (b) Channel, (c) Tube, (d) Cylinder, (e) Open path with core


The results of the simulations can be found here. From our modelling report, we can conclude that our results concur with the results of Wong et al. (2008), the open path with core was found to be the most optimal. This was determined by conducting a holistic comparison of SVM stress, SVM strain, displacement and the applied force on each scaffold.

Learn: Do the solutions meet the requirements?

After testing the initial designs and deciding on the open path with core design we were apprehensive about our results. The main concern was that the simulations may have been too simple with only the gravitational load being applied. We arranged a meeting with Dr. Novak Elliot: Adjunct lecturer in the Department of Mechanical Engineering at Curtin University. Dr. Elliot advised us on how to add further parameters to our model. He suggested simulating a cough inside the patient, of which there would be a pressure gradient along the spinal cord, thus exerting a pressure load on the scaffold.

Improve: Based on the results and data, make design changes, prototype, test again, and review new data

From Martin & Loth (2009), we obtained values for transmural pressure. Transmural pressure is the pressure difference across a hollow structure; it is the pressure gradient across the vessel wall. The compression of the syringomyelia is caused by the transmural pressure force. The results are displayed here. They show that when considering gravitational load and the simulation of a cough, the scaffold is capable of withstanding the force and will not break. Furthermore, we segmented anonymous patient scans using ITK SNAP and segmented them to obtain the shape of the cyst. With this information, we were able to make suitable adjustments to our scaffold dimensions to ensure it would fit our patient. You can watch a video tutorial on how to do this on our contribution page here.

Micro-architecture is defined to be the surface structure or topography of the scaffold and consists of features including, but not limited to, porosity (alongside mean pore size and shape), microgrooves and micro-channels. It is known that these microscale factors have a significant effect on cell behaviour and differentiation (Meco and Lampe, 2018). Furthermore, the axonal response to the implant is subject to variation based upon substrate topography and the size of the aforementioned structures (Zhang et al., 2020). Consequently, this aspect of the design is imperative to ensure the success of the scaffold. Two of the initial topographies that we planned to implement comprised of microgrooves and pores, as these were the most promising within the literature.


Physical guidance during nerve regeneration is of utmost importance, and microgrooves have been shown to be an effective tool to achieve such with respect to neurites in vitro (Goldner et al., 2006). The contacts between the grooved surface and axons allow for guided elongation along groove edges—essentially forming bridges of neurites (Zhang et al., 2020).

Figure 2: Microgroove microstructure, along with axonal growth (Goldner et al., 2006)

However, the optimisation of groove design is difficult because the dimensions of the grooves play a large role in differentiation and alignment—and the axonal response to the groove size is known to be dependent on the specific neuron (Meco and Lampe, 2018). Consequently, selecting a uniform size for the grooves proves to be challenging due to the potential injury site having a variety of cell and neuron type. Furthermore, the maximum limit for groove plateau width to support neurite alignment in vitro lies within the range of 10-30μm (Wong et al., 2008). Conversely, this raises further issues because our proposed method of scaffold production is 3D bioprinting; common printer types including FDM (Fusion Deposition Modelling) (Salentijn et al., 2017) and SLA (Stereolithography Apparatus) (Lee, Ng and Yeong, 2019) have a resolution limit of approximately 100μm. Therefore, the addition of microgrooves would add another stage in production, such as soft lithography (Sun, Ferrell and Hansford, 2004)—making it significantly more complicated, potentially less repeatable and less time effective. Furthermore, this extra production stage would lead to a higher cost, and hence our solution would become less affordable. After conversing with Dr. Jacob Koffler, a subject-matter expert who advised us against the over-complication of the scaffold micro-architecture, we decided to no longer pursue this avenue of design.


Our chosen scaffold material, polycaprolactone, has many strengths; nevertheless, it does not possess all of the properties required to be an optimally functional scaffold for spinal cord injury, due to its unsuitable stiffness and cell adhesion in its dense form (Shahriari et al., 2017). A solution for this, detailed by Shahriari et al. (2017) and Guarino, Causa and Ambrosio (2007), is to introduce porosity.

The Young’s Modulus of the spine is around 3.5 kPa (Ozawa et al., 2001), as opposed to the 300MPa modulus of dense PCL (Eshraghi and Das, 2010)—a difference in magnitude of approximately 105. Although, with the incorporation of pores this may be reduced to 0.24MPa at a porosity of 91% (Guarino, Causa and Ambrosio, 2007), a more acceptable difference of 102 in order. A further variable explored throughout by Guarino, Causa and Ambrosio (2007) was pore size, and its effect upon the elastic modulus; the findings presented showed that this factor had less of a significant effect than the overall porosity percentage (or open volume)—therefore we reached the conclusion that the porosity percentage should be used to determine the elastic properties, whereas the size of the pores should be determined by the optimal value for other pore functions. Currently, the accepted range for porosity within scaffold engineering is > 50% (Bayram et al., 2019). In spite of this, further experimentation has shown that polycaprolactone scaffolds become too brittle and risk fracture at percentages above 70% (Shahriari et al., 2017) – resulting in a more realistic range of 50-70% for ideal porosity. Considering that the scaffold should have a fine balance between a higher open pore volume for growth of axons (Thomas et al., 2013), and still be mechanically sound, we decided on a porosity percentage of 60%, which has been shown to have an elastic modulus of around 1.48MPa (Guarino, Causa and Ambrosio, 2007).

Interconnected porous networks are essential for cell nutrition, proliferation and migration for the formation of new tissues (Loh and Choong, 2013). Typically, the accepted ideal pore size is within the range of 100 to 400μm (Bayram et al., 2019)—however, this is a very large range to select a specific pore size from. Some studies have stated that small pores and a lack of interconnectivity can lead to cell necrosis (Germain et al., 2018); therefore, we decided that we should select a value towards the middle of this range. An experiment carried out by Qing-song et al. (2010), suggested a pore size of 200-300μm for creating spinal cord scaffolds – and so we decided to design our scaffold based upon this.

A feature that is closely tied to pore size is pore shape. Within computer-aided design (CAD) and its associated computational modelling, the overall scaffold can be broken down into voxels (uniformly sized volumetric cubes), of which can be filled by a cell—the shape of such can be determined by the user. These cells then become the constituents of the pores within the scaffold. We chose to implement gyroid morphology owing to its high interconnectivity, good nutrient and waste diffusion, and because scaffolds made up of a gyroid structure tend to retain their mechanical robustness throughout the degradation stages (Germain et al., 2018). Furthermore, gyroid structures have been shown to be more favourable for cell adsorption, and have the greatest number of micropores (within the magnitude of 10μm) introduced throughout degradation - as opposed to other topology structures such as spherical pores (Jin et al., 2019). This is of great benefit because micropores have been shown to improve cell attachment properties, hence making the scaffold more suitable for nerve guidance (Shahriari et al., 2017).

Figure 3: Gyroid structure (left), alongside singular gyroid cell (right) (Abueidda et al., 2019)

Computational Design and Modelling

3D bioprinting allows for the structure of the scaffold to be designed with great accuracy. Not only does this permit repeatable and reproducible results but allows a predetermined and specific pore distribution. Further literature studies have found that a heterogeneous distribution of pore size and placement results in a variation of mechanical properties throughout the scaffold (Bayram et al., 2019), as well as inhibiting cellular activities and preventing the formation of homogeneous tissues (Choi, Zhang and Xia, 2010). Hence, regularity within the scaffold design became a large focus in terms of research.

The macrostructure of the scaffold (open path with core) was designed within Autodesk Inventor; however, the complexity of the microstructure design could not be implemented with this software. Therefore, we imported the scaffold design from Inventor into Rhinoceros by exporting the macro design in .dwg format. Subsequently, the Grasshopper visual coding language was utilised to manipulate the structure within Rhinoceros, and the steps behind the algorithm for this process to create the pores can be seen in the following flow chart. We called this algorithm 'The Porosifier'.

Final Porous Design
Figure 4: Original macro-architecture of scaffold (left) and porous scaffold design (right)

Table 2: Porous scaffold properties generated within Rhinoceros using Grasshopper

Unit Cell Size 0.4mm
Mesh Thickness 0.1mm
Original Scaffold Volume 166.6mm3
Porous Scaffold Volume 70.44mm3
Average Pore Size 199.5μm
Porosity 58%

The final porous scaffold design can be seen within Fig. 4, with the scaffold properties as shown within Table 2. The average pore size generated using the specific parameters detailed was approximately 200μm, which is within the accepted range as specified in the aforementioned literature. The porosity of the scaffold is also within the accepted range yet is 2% lower than our desired 60% - this is because it was difficult to program the exact parameters, especially because the software is computationally expensive. This porosity was calculated using the following equation, as found by (Loh and Choong, 2013):

\[Porosity = 1 - \frac{V_{lattice}}{V_{total}} \times 100\]

Where V represents volume.

These properties for porosity and pore size were found by modulating the unit cell size and thickness until the desired approximate values were reached. This method of designing the pores was guided by Aaron Porterfield (founder of F=F), whom designed the main plug-in we used for the lattice generation – Crystallon. We have made The Porosifier open-source, and it may be applied to any arbitrary scaffold or shape, and is easily adjustable. The grasshopper file for this can be found on our GitHub, here


Scaffold Requirements

  • Sakiyama-Elbert, S., Johnson, P., Hodgetts, S., Plant, G. and Harvey, A., 2012. Scaffolds to promote spinal cord regeneration. Handbook of Clinical Neurology, pp.575-594.
  • Tabesh, H., Amoabediny, G., Nik, N., Heydari, M., Yosefifard, M., Siadat, S. and Mottaghy, K., 2009. The role of biodegradable engineered scaffolds seeded with Schwann cells for spinal cord regeneration. Neurochemistry International, 54(2), pp.73-83.
  • Wang, Y., Tan, H. and Hui, X., 2018. Biomaterial Scaffolds in Regenerative Therapy of the Central Nervous System. BioMed Research International, 2018, pp.1-19.
  • Wong, D., Leveque, J., Brumblay, H., Krebsbach, P., Hollister, S. and LaMarca, F., 2008. Macro-Architectures in Spinal Cord Scaffold Implants Influence Regeneration. Journal of Neurotrauma, 25(8), pp.1027-1037.

Material Choice

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  • Dong, C. & Lv, Y. 2016. Application of Collagen Scaffold in Tissue Engineering: Recent Advances and New Perspectives. Polymers (Basel), 8.
  • Du, B. L., Zeng, C. G., Zhang, W., Quan, D. P., Ling, E. A. & Zeng, Y. S. 2014. A comparative study of gelatin sponge scaffolds and PLGA scaffolds transplanted to completely transected spinal cord of rat. J Biomed Mater Res A, 102, 1715-25.
  • Madihally, S. V. & Matthew, H. W. 1999. Porous chitosan scaffolds for tissue engineering. Biomaterials, 20, 1133-42.
  • Sakiyama-elbert, S., Johnson, P. J., Hodgetts, S. I., Plant, G. W. & Harvey, A. R. 2012. Scaffolds to promote spinal cord regeneration. Handb Clin Neurol, 109, 575-94.
  • Siddiqui, N., Asawa, S., Birru, B., Baadhe, R. & Rao, S. 2018. PCL-Based Composite Scaffold Matrices for Tissue Engineering Applications. Mol Biotechnol, 60, 506-532.
  • Zarrintaj, P., Manouchehri, S., Ahmadi, Z., Saeb, M. R., Urbanska, A. M., Kaplan, D. L. & Mozafari, M. 2018. Agarose-based biomaterials for tissue engineering. Carbohydr Polym, 187, 66-84.
  • Zhang, L., Fang, H., Zhang, K. & Yin, J. 2018. Homologous Sodium Alginate/Chitosan-Based Scaffolds, but Contrasting Effect on Stem Cell Shape and Osteogenesis. ACS Appl Mater Interfaces, 10, 6930-6941.
  • Zhang, L., Yang, G., Johnson, B. N. & Jia, X. 2019. Three-dimensional (3D) printed scaffold and material selection for bone repair. Acta Biomater, 84, 16-33.


  • Budday, S., Nay, R., de Rooij, R., Steinmann, P., Wyrobek, T., Ovaert, T. C., & Kuhl, E. (2015). Mechanical properties of gray and white matter brain tissue by indentation. Journal of the Mechanical Behavior of Biomedical Materials, 46, 318–330.


  • Abueidda, D. W. et al. (2019) ‘Mechanical properties of 3D printed polymeric Gyroid cellular structures: Experimental and finite element study’, Materials & Design, 165, p. 107597. doi: 10.1016/j.matdes.2019.107597.
  • Bayram, C. et al. (2019) ‘Biofabrication of Gelatin Tissue Scaffolds with Uniform Pore Size via Microbubble Assembly’, Macromolecular Materials and Engineering, 304(11), p. 1900394. doi: 10.1002/mame.201900394.
  • Choi, S.-W., Zhang, Y. and Xia, Y. (2010) ‘Three-Dimensional Scaffolds for Tissue Engineering: The Importance of Uniformity in Pore Size and Structure’, Langmuir, 26(24), pp. 19001–19006. doi: 10.1021/la104206h.
  • Eshraghi, S. and Das, S. (2010) ‘Mechanical and microstructural properties of polycaprolactone scaffolds with one-dimensional, two-dimensional, and three-dimensional orthogonally oriented porous architectures produced by selective laser sintering’, Acta Biomaterialia, 6(7), pp. 2467–2476. doi: 10.1016/j.actbio.2010.02.002.
  • Germain, L. et al. (2018) ‘3D-printed biodegradable gyroid scaffolds for tissue engineering applications’, Materials & Design, 151, pp. 113–122. doi: 10.1016/j.matdes.2018.04.037.
  • Goldner, J. S. et al. (2006) ‘Neurite bridging across micropatterned grooves’, Biomaterials, 27(3), pp. 460–472. doi: 10.1016/j.biomaterials.2005.06.035.
  • Guarino, V., Causa, F. and Ambrosio, L. (2007) ‘Porosity and mechanical properties relationship in PCL porous scaffolds’, Journal of applied biomaterials & biomechanics: JABB, 5(3), pp. 149–157.

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